Methods for treating dental conditions using tissue scaffolds

ABSTRACT

The invention provides methods, apparatus and kits for regenerating dental tissue in vivo that are useful for treating a variety of dental conditions, exemplified by treatment of caries. The invention uses tissue scaffold wafers, preferably made of PGA, PLLA, PDLLA or PLGA dimensioned to fit into a hole of corresponding sized drilled into the tooth of subject to expose dental pulp in vivo. In certain embodiments the tissue scaffold wafer further comprises calcium phosphate and fluoride. The tissue scaffold wafer may be secured into the hole with a hydrogel, a cement or other suitable material. Either the wafer or the hydrogel or both contain a morphogenic agent, such as a member encoded by the TGF-β supergene family, that promotes regeneration and differentiation of healthy dental tissue in vivo, which in turn leads to remineralization of dentin and enamel. The tissue scaffold may further include an antibiotic or anti-inflammatory agent.

This application is a Continuation claiming benefit under 35 U.S.C. §120 of U.S. patent application Ser. No. 10/684,226, filed Oct. 10, 2003,which is hereby incorporated by reference in its entirety.

TECHNICAL FIELD

This invention relates generally to the field of treating dentalconditions, particularly caries, and more particularly to methods,compositions and devices that promote in vivo regeneration andremineralization of dentin and enamel by inserting tissue scaffoldmaterials in vivo, into holes drilled into a tooth having need of dentinregeneration.

BACKGROUND OF THE INVENTION

The development of tissue scaffold materials for regenerating tissue exvivo and for uses of such ex vivo regenerated tissue/scaffoldcombinations to treat patients in vivo has been a growing subject ofinterest in the prior art. Of relevance to the invention describedhereinafter are prior art uses of tissue scaffolds for ex vivo cultureof oral tissues.

U.S. Pat. No. 5,885,829 discloses use of tissue scaffolds made of aporous matrix for the ex vivo culturing and regeneration of oral tissuesfrom isolated dental cells. Numerous polymers, both biodegradable andnon-biodegradable, synthetic or natural scaffold materials weredescribed as suitable for culturing oral tissues ex vivo. In particular,homopolymers of poly lactic acid (PLLA), poly[D,L-lactic acid] (PDLLA),poly-glycolic acid (PGA) and heteropolymers of lactic acid and glycolicacid, i.e., poly[lactic-co-glycolic acid] (PLGA), alone or incombination with polyvinyl alcohol (PVA), were shown to be effective forculturing fibroblasts isolated from dental pulp. Cells from dental pulpwere first explanted, separated and propagated in a monolayer cultureusing ordinarily tissue culture techniques. In one exemplified method,the cultured cells were removed and then seeded onto a matrix of about 3mm thickness made of a mat of PGA fibers. The seeded matrix was thenincubated in growth medium and it was shown that the culturedfibroblasts were able to adhere to the PGA fibers and ultimately occupythe spaces between fibers. Tubular matrix scaffolds were made fromporous PLLA, PDLLA and PLGA two dimensional films using a particulateleaching technique. A three dimensional tubular matrix device wasconstructed by stacking the films on one another and chemically sealingthe edges. The tubular matrix device was shown to allow growth ofvascular tissue when implanted into the mesentery and omentum of rats.Of the combinations tested, PLLA and PDLLA were shown to maintain theirstructure in vivo, while PLGA based devices did not. PGLA devices weredetermined to have less resistance against compressional forces thanPLLA or PDLLA. Other techniques for forming tubular matrices were alsodescribed, including bonding of PLLA, PGA, and PLGA tubes using achloroform spray. In another example, a three dimensional “spongematrix” was described for a tissue scaffold formed of PLA infiltratedwith PVA or of PLGA at a ratio of 85:15 D,L lactic acid:glycolic acid.These sponge devices were shown to support growth of seeded liver cellsex vivo. The devices were also implanted into the mesentery of rats withand without seeded cells and shown to support in growth of vasculartissue in vivo.

Human dental pulp cells were shown to propagate ex vivo better on a PGAmatrix than on collagen or alginate based hydrogel matrices. Seededcells grown in this manner ex vivo filled the spaces within the PGAmatrix, which was eventually replaced by new tissue including tissuecontaining collagen indicating formation of an extracellular matrix. Inanother experiment, human gingival cells and pulp derived fibroblastswere shown to infiltrate a PGA matrix when seeded and cultured ex vivo,and to express human gene products when the seeded and cultured matrixwas implanted subcutaneous in mice in vivo, even though the majority ofcells in the implant were mouse fibroblasts. In addition it wasdisclosed that tissue scaffold materials could be used to deliver drugsby demonstrating that epidermal growth factor (EGF) could be entrappedin microspheres made of a 75:25 PLGA copolymer and slowly released intoa buffer in vitro over 30 days. Further, hepatocyte cells seeded intocylindrical microspheres containing EGF and implanted into the mesenteryof rats were shown to exhibit the biological effects of EGF onhepatocyte cellular activity over a 4 day period when implanted in vivo.Although this patent discloses a utility for using scaffold material forex vivo propagation of oral tissues, it fails to disclose anytherapeutic methods for the use of such ex vivo cultured cells fortreating teeth in vivo, and fails to disclose the use of tissuescaffolds in the absence of ex vivo culturing.

U.S. Pat. No. 6,281,256 discloses a process of preparing open porematrices of a biodegradable polymer made of PLLA, PGA or PGLA polymersby using gas forming and particulate leaching steps to form pores in thematrices of the polymers. In a typical example of the disclosed process,a PLGA copolymer is formed in a mixture that includes a leachableparticulate material. The mixture is molded, optionally undercompression, to a desired size and shape and is subject to a highpressure gas atmosphere to dissolve gas in the copolymer. Then athermodynamic instability is created by reduction of the pressure sothat the dissolved gas nucleates and forms gas pores within thecopolymer, causing expansion and fusion of the copolymer particles,creating continuous polymeric matrix still containing the particulatematerial. The particulate material is then leached from the polymericmatrix with a leaching agent creating a porous matrix. The amount andsize of the particulate material used in the mixture determines thelevel of interconnectivity between open pores, the size of the pores andthe amount of pores in the final matrix (porosity). The interconnectedmatrices formed by this method have a porosity of between about 25% to95-97% and exhibit high tensile strength with a tensile modulus of about850 to 1100 kPa and a compression modulus of about 250 kPa or larger.Such matrices were disclosed as being useful for bone formation andguided tissue regeneration (GTR) where tissues could be grown within thematrix pores guided by the scaffolding material, which provides asurface for cellular attachment. The porous matrices could also be madeto have a non-porous barrier one on end by forming an impermeable skin,or could be made of different levels of porosity throughout by alteringthe amount of leachable particulate material in different sections sothat one section forms open pores and another does not. Such matriceswhere also shown to be capable of releasing a growth factor VEGF invitro, over a period of 20 to 21 days. Use of such matrices wasgenerally mentioned as having utility in regenerating oral tissues. Theuse of such matrices configured with an impermeable side wasparticularly suggested for treating periodontal disease by using thepores in one section of the matrix to grow periodontal ligament cellsand providing a barrier in another surface of the matrix to prevent downgrowth of epithelial cells. However, no actual method was described forthe use of such matrices in vivo without previously seeding the matrixand culturing the cells ex-vivo.

U.S. provisional patent application No. 60/166,191 describes methods forproducing tissue scaffolds of PGLA using fused salt crystals of selectedsizes to control the porosity, interconnectivity and ease of manufactureof the scaffolds.

U.S. Pat. No. 6,472,210 discloses a method of making a polymer scaffoldhaving an interconnected passage-way of strutted pores with diameters inthe range of about 0.5 to 3.5 mm. The polymer may be PLGA. The polymeris prepared by mixing a liquid polymer in a solution with an organicsolvent such as DMSO, methylene chloride, ethyl acetated chloroform,acetone, benzene butanone, carbon tetrachloride, heptane, hexane orpentane. The liquid polymer solution is mixed with particles of 0.5 toabout 3.5 mm in diameter. The particle/polymer mixture is then treatedwith a “non solvent” for the polymer, such as water, alcohol, dioxane oraniline, in a phase inversion step that precipitates thepolymer/particle. The particle is then leached from the precipitate bytreatment with a solvent that dissolves the particle material but notthe polymer. The method and composition are said to be suitable forforming tissue scaffolds for use in regenerating tabecular bone, whichhas a high porosity and large strutted trabeculae widths on the order ofabout 0.14 mm to about 0.3 mm. Such large macrospores would not besuitable for regenerating dentin or other oral tissues in the teeth.

WO 00/56375 and Murphy and Mooney, (2002) J. American Chemical Society124(9) 1910-1917, disclose methods for patterning (mineralizing) tissuescaffold material formed into three dimensional wafers for bioimplants.In one aspect, the surface of tissue scaffolds such as PGA, PLA and PLGAare coated with minerals such as calcium chloride, and phosphate usefulfor orthopedic tissue mineralization. The scaffold material is treatedby electromagnetic radiation or by an electron beam to cause surfacedegradation via photolysis or electrolysis. Lithographic techniques aredisclosed for forming patterns on the surface to create the desiredsites of degradation. Alternatively, chemical hydrolysis of the surfaceor direct soaking of the scaffold material in an appropriate mineralsolution may be used to pattern the wafer. In any case, the modifiedsurface of the scaffold contains functional groups, such as polar oxygengroups (carboxylates in particular) that promote calcium phosphateformation on the surfaces of the materials used to form the scaffoldmaterials when the treated scaffold is immersed in an appropriatesolution. Osteogenic cell precursors may be seeded onto the mineralizedbiomaterial ex vivo. Alternatively, bone cells were said to attach tothe mineralized scaffold material in vivo. The growth factor VEGF wasshown to be released from such mineralized tissue scaffold materialsover time. While numerous utilities of this patterning technique aredisclosed, the patent fails to teach any therapeutic method that usesthe patterned tissue scaffolds for a therapeutic treatment of dentalconditions in vivo.

Other publications describe use of tissue scaffold material or hydrogelsto deliver morphogenic agents (or genes encoding the same) that promotegrowth or development of various tissues in vivo after ex vivo culture.Rutherford, R. B., (2001) Euro. J. Oral Science 109(6) 422-444 disclosedthat ex vivo grown dermal fibroblasts transduced with an adenovirusexpression vector expressing a cDNA encoding bone morphogenic protein 7(BMP-7) were effective at inducing reparative dentinogenesis withapparent regeneration of the dentin-pulp complex when transplanted invivo in ferrets having pulpitis. However, no effect was seen when a 10fold range (2.5 μg to 25 μg recombinant protein) was delivered directlyto the pulp tissue. Rutherford R. B., et al (2000) Eur. J. of OralScience 108(3) 202-206.

Sloan et. al. (2000) Archives of Oral Biology 45(2) 173-177 used ahydrogel of agarose beads soaked in BMP-7 to deliver the protein totooth slices cultured ex vivo in a semi solid agar and disclosed alocalized increase in extracellular matrix secretion by odontoblasts atthe site of application. Nakashima, M., (1994) Archives of Oral Biology(12) 1085-89 demonstrated that recombinant BMP-2 and BMP-4 induceddentin formation in amputated pulp of dogs in vivo when condensed on apowdered carrier comprised of dried type 1 collagen and proteoglycans.In a prior publication, BMP-2 and BMP-4 were shown to induce dentinformation in amputated pulp when condensed on the same the same type ofdelivery material. Nakashima, M., (1994) J. Dental Research 73(9)1515-1522.

Other types of materials such as various modified hydrogels also providetissue scaffold-like functions for propagating various tissue. Andersonet al, (2002) PNAS 17; 99(19), 12025-30, disclosed that an alginatebased hydrogel modified with the tripeptide sequence RGD promoted cellmultiplication and bone tissue-like growth plates when chondrocytes weretransplanted ex vivo.

U.S. Pat. No. 6,413,498 discloses a mixture of cationic and anionic ionexchange resins charged with Ca²⁺, F⁻ and PO₄ ³⁻ in molar ration of2:1:1 for use in a filler for the treatment of caries. These resins,typically made of polystyrene, promote remineralization of dentin toform tissue having a composition and hardness close to that of originaldentin. Such resins are disclosed to also be useful as components ofdentifrice products are not suitable as a tissue scaffold to promoteregeneration of the cell types that are required for healthy dentin.Moreover, such resins are believed to leave organic residue upon contactwith teeth.

U.S. Pat. Publication No. 2002/0119180 A1 and Young et al., (2002) J.Dent. Res. 81 [10] p 695-700 each describe regenerating multiple dentaltissues in organized tooth structures in situ by ex vivo seeding ofenamel and pulp organ tissues on a PGA/PLLA or PLGA scaffold formed inthe shape of [[a]] human teeth. The ex vivo cultured tissue/scaffoldcombinations were collagen coated and then implanted in the omentum ofrats where they were cultured in situ. The in situ cultured tissue wereshown to develop mineralized structures indicative of enamel surroundingdentin, and to develop into odontogenic cell types, including,ameloblasts, odontoblast-like cells, putative cemetoblasts and cementum.While the experiment demonstrated the potential feasibility ofregenerating whole teeth by ex vivo seeding and in vivo culturing ofisolated dental tissue, Young et al, however, did not disclose anymethod of treating dental tissue in vivo.

While the prior art recognizes the utility of using tissue scaffolds forgrowing tissue in vitro or for treating bone lesions in vivo, thereremains a need in the art for methods and devices for treating dentalconditions in vivo using such tissue scaffold. The present inventionprovides for such methods and devices.

SUMMARY OF THE INVENTION

The present invention provides methods, compositions and devices basedon the discovery that tissue scaffolding materials can be directly usedto facilitate regeneration of dentin in a tooth of subject in vivo,without need of seeding of the scaffold material by ex vivo culture ofcells prior to implanting the scaffold material into dental tissue. Inthe methods of the invention, the scaffolding material is used as asubstitute or as an enhancement of prior art methods for treatingvarious stages of dental caries or pulpitis. The methods are suitablefor treating conditions ranging from asymptomatic caries where only asmall portion of dentin below the crown of the tooth is degenerated, totreating deep caries where dentin is degenerated down to the rootordinarily requiring a root canal treatment by the methods of the priorart. The methods apply the scaffolding material directly intoappropriately sized holes drilled into a subject's teeth so that atleast a portion of the pulp is exposed. The scaffolding material isimplanted into the hole so that a portion of the scaffold material is incontact with the exposed portion of the pulp. Pulp cells are stimulatedto grow into the matrix of the scaffolding material causingremineralization and formation of new dentin.

More particularly, one aspect of the invention is a method for treatinga subject's tooth in need of regeneration of dentin that includes theacts of forming a hole in the tooth of the subject in vivo, the holebeing of a depth sufficient to expose at least a portion of pulp,inserting a tissue scaffold into the hole so that a portion of thetissue scaffold contacts at least a portion of the exposed pulp; andregenerating dentin by allowing sufficient time for tissue to grow invivo, from the pulp into the tissue scaffold and to regenerate thedentin. In certain embodiments, the invention the tissue scaffoldinserted into the hole does not include an ex vivo cultured tissuewithin the scaffold. In other embodiments, dental pulp stem cells areseeded into the tissue scaffold and cultured therein prior to insertionof the tissue scaffold into the hole. In still other embodiments, dentalpulp stem cells are added directly to the region of the exposed pulpprior to inserting the tissue scaffold into the hole.

In certain embodiments, the tissue scaffold is formed into a shapedimensioned to fit snuggly into the hole that is formed so that thetissue scaffold does not move more than 0.1 mm in a lateral direction inthe hole. The tissue scaffold is typically formed into a cylindricalwafer having a diameter of about 2 to about 5 mm and a height of about0.5 to about 2 mm.

In various embodiments, the tissue scaffold also contains calciumphosphate associated therewith. In particular embodiments the tissuescaffold also contains fluoride associated therewith.

In certain embodiments, the tissue scaffold is comprised of scaffoldingmaterial selected from the group consisting of PLA, PGA, PDLLA PLLA andPLGA. In a particular embodiment, the tissue scaffold is comprised ofscaffolding material is comprised of PLGA.

In certain embodiments, the tissue scaffold may further include aphysiologically effective amount of a morphogenic agent that promotesgrowth of dentin tissue or the mineralization thereof. In particularembodiments, the morphogenic agent is selected from a protein encoded bya TGF-β supergene family. In more particular embodiments, the protein isselected from the group consisting of BMP-2, BMP 4, BMP-7, VEGF, FGF-1,FGF-2, 1GF-1, 1GF-2, PDGF, GDF-1, GDF-2, GDF-2, GDF-3, GDF-4, GDF-5, orcombinations of the same. In still more particular embodiments theprotein is selected from the group consisting of BMP-2, BMP 4, BMP-7,and GDF-5. In yet another more particular embodiment, the morphogenicagents includes at least one of PDGF VEGF, and a protein selected fromthe group consisting of BMP-2, BMP 4, BMP-7, and GDF-5.

In certain embodiments, the tissue scaffold further includes an activeagent selected from the group consisting of an anti-bacterial agent, ananti-inflammatory agent, and an analgesic agent. Example antibioticsinclude antibacterial agents such as tetracycline. Exampleanti-inflammatory agents include COX I and II inhibitors. Exampleanalgesic agents include anti-inflammatory and anaesthetic agents.

In certain embodiments, the methods of the invention may includecovering the hole with a cement. In particular embodiments the cementmay be comprised of a combination of di-calcium and tetra-calciumphosphate. In certain other embodiments, the cement or amalgam iscomprised of calcium phosphate and fluoride.

In one embodiment of the methods of the invention, the subject hasasymptomatic caries and the act of forming the hole exposes a portion ofthe pulp located in a coronal cap region of the tooth and the tissuescaffold is inserted to contact the exposed pulp. In another embodiment,the subject has need of a pulpotomy and the act of forming the holeexposes a portion of the pulp located in a coronal region of the toothand the tissue scaffold is inserted into the coronal region to contactthe exposed pulp. In another embodiment, the subject has need of a rootcanal and the act forming the holes exposes a portion of the pulp in atleast one of the root canal, and a coronal region of the tooth and thetissue scaffold is inserted into the hole to contact the exposed pulp.

In another aspect, the invention includes devices for treating a tooththat include a tissue scaffold comprised of a scaffolding polymerconfigured as a wafer that fits snuggly into a corresponding hole thatis formed in a tooth of the subject so that the tissue scaffold does notmove more than 0.1 mm in a lateral direction in the hole. In certainembodiments, the hole of corresponding size is formed by an act ofdrilling the tooth. In some embodiments the wafer is cylindrical and hasa diameter of about 2 to about 5 mm and has a height of about (0.1 toabout 0.5 mm.

In yet another aspect, the invention provides compositions for treatingdental tissue that include: a tissue scaffold comprising a scaffoldingmaterial associated with calcium phosphate and fluoride. The compositionmay further comprise a physiologically effective amount of a morphogenicagent that promotes growth of dentin tissue. In particular embodiments,the morphogenic agent is selected from a protein encoded by a TGF-betasupergene family. In more particular embodiments, the protein isselected from the group consisting of BMP-2, BMP 4, BMP-7, VEGF, FGF-1,FGF-2, IGF-1, IGF-2, PDGF, GDF-1, GDF-2, GDF-2, GDF-3, GDF-4, GDF-5, orcombinations of the same. In still more particular embodiments theprotein is selected from the group consisting of BMP-2, BMP 4, BMP-7,and GDF-5. In yet another more particular embodiment, the morphogenicagents includes at least one of PDGF VEGF, and a protein selected fromthe group consisting of BMP-2, BMP 4, BMP-7, and GDF-5. The compositionmay further comprise an active agent selected from the group consistingof an anti-bacterial agent, an anti-inflammatory agent, and an analgesicagent. The composition may include an active agent selected from thegroup consisting of an anti-bacterial agent, an anti-inflammatory agent,and an analgesic agent. In typical embodiments, the composition iscomprised of scaffolding polymer selected from the group consisting ofPLA, PGA, PDLLA PLLA and PLGA. In preferred embodiments the scaffoldingpolymer is PLGA.

In some embodiments the device is made of a scaffolding materialselected from the group consisting of PLA, PGA, PDLLA PLLA and PLGA. Insome embodiments the scaffolding material is associated with calciumphosphate. In some embodiments the wafer is associated with calciumphosphate and fluoride. The device may further include apharmaceutically acceptable carrier, buffer, excipient or diluent.

In some embodiments, the device has a top surface, a bottom surface anda side perimeter surface between the top and bottom surfaces, and atleast one of the top and bottom surfaces are marked with a pattern thatalters appearance when the wafer is crushed. In one embodiment thepattern is comprised of set of concentric circles. In anotherembodiment, the pattern is comprised of a dye that alters color when thewafer is compressed.

In another aspect, the invention provides a kit containing a pluralityof the forgoing wafers wherein the plurality of wafers are of aplurality of sizes selected to correspond with a plurality of holesizes. The plurality of hole sizes correspond to a plurality of holesmade by drilling a tooth with any one of a plurality of dental drill bitsizes. In one embodiment, the kit includes, a dry tissue scaffold waferdimensioned to be received into a hole of corresponding size formed in atooth of a subject and a well configured to hold the tissue scaffoldwafer in a dry state. Also included is a second well adjacent to thefirst well, the second well holding a hydrating liquid comprising apharmaceutically acceptable liquid. A breakable partition separates thefirst and second wells, the breakable partition being structured tobreak by a force applied by a human hand causing the tissue scaffoldwafer to contact hydrating liquid when the breakable partition isbroken.

In yet another aspect, the invention provides a support casing toprotect tissue scaffold from crushing. The casing a least partiallysurrounds the tissue scaffold wafer. The tissue scaffold wafer is madeof a porous scaffolding material having a first crushing resistance andthe support casing is made of a material having a second crushingresistance greater than the first crushing resistance. In a typicalembodiment, the support casing includes at least one horizontallydisposed member that contacts at least one of an upper and lower surfaceof the tissue scaffold wafer and includes at least one columnarextension extending from the horizontally disposed member along a sideperimeter of the tissue scaffold wafer.

Yet another aspect of the invention provides a vacuum manipulator formanipulating a tissue scaffold wafer. The vacuum manipulator includes avacuum tube having a proximal end, a distal end, and walls between theproximal and distal ends enclosing the vacuum tube. The manipulatorincludes an attachment for a vacuum source in fluid communication withthe vacuum tube; and a suction cup attached to the proximal end of thevacuum tube in fluid communication with the vacuum tube. The suction cupis sized to fit onto a surface of a wafer comprised of dentalscaffolding material. The manipulator also includes a valve assembly atthe distal end of the vacuum tube, which is operable to close and openfluid access between the vacuum tube and a vacuum source and to open andclose fluid access between the vacuum tube and a pressure source.

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1 illustrates tooth anatomy for reference to various aspects of theinvention.

FIG. 2 illustrates an example of a tissue scaffold wafer suitable foruse in one aspect of the invention.

FIGS. 3A-3D illustrate various markings on the top surface of wafersaccording to another aspect of the invention. FIG. 3A shows concentricmarkings. FIG. 3B shows perpendicular hatch markings. FIG. 3C showsparallel line markings. FIG. 3D shows dot markings.

FIGS. 4A-4D illustrate various tissue scaffold wafer casing supportsaccording to another aspect of the invention. FIG. 4A shows a ringsupport. FIG. 4B shows a pad support. FIG. 4C shows a bracket support.

FIG. 4D shows a brace support.

FIG. 5 illustrates a tissue scaffold wafer with casing supports and ahydrogel seal implanted in a tooth according to another aspect of theinvention.

FIG. 6 illustrates a tissue scaffold wafer kit according to anotheraspect of the invention.

FIG. 7 illustrates a vacuum manipulator according to another aspect ofthe invention.

FIG. 8 illustrates treating asymptomatic caries according to anotheraspect of the invention.

FIG. 9 further illustrates asymptomatic treating caries according to theinvention.

FIG. 10 illustrates treating coronal pulp carries according to anotheraspect of the invention.

FIG. 11 illustrates a tooth treated with a root canal procedureaccording to the prior art.

FIG. 12 illustrates a replacement for a root canal according to anotheraspect of the invention.

DETAILED DESCRIPTION OF THE INVENTION

In the description that follows, citation is made to numerous referencesthat may aid one of ordinary skill in the art to understand or practiceof the invention in its fullest scope. Each such reference, as well asthe references cited therein, is incorporated herein by reference to theextent needed to practice the methods or make the materials of thepresent invention, and to the extent such references do not conflictwith the teachings of this invention, in which case the teachings ofthis invention are to be used to substitute or supplement suchconflicting teachings in the incorporated references.

Abbreviations and Definitions

Prior to setting forth the invention in detail, certain terms aredefined herein below as better aid in an understanding thereof. Thedefinitions below are not intended to limit the ordinary meaning of theterms as they would be understood by one of ordinary skill in the art,except if the ordinary meanings would conflict with the meanings setforth below. Otherwise, the definitions are intended to illustratevarious aspects of the ordinarily understood meanings:

“Biodegradable” with respect to a composition or device means thecomposition or device may be eroded, absorbed, digested, destroyed,reconfigured, removed or otherwise degraded from its original form whenplaced into contact with a selected animal tissue or fluid in vivo for aperiod of 30 days or less. The animal tissue or fluid may or may notinclude flora such as bacteria, fungi or other microbe that facilitatesthe degradation. The typical tissue or fluid involved in biodegradationincludes any oral or dental tissue or part thereof. If the compositionor device remains substantially in its original form after 30 days of invivo contact with animal tissue, it is herein considered“non-biodegradable.”

A “tissue scaffold” is any composition formed into a porous matrix intowhich tissue can grow in three dimensions. Tissue scaffolds of thepresent invention are at least 60% porous, at least 70% porous, at least80% porous, at least 90% porous, or at least 97% porous, where porosityis a measure of the volume of liquid, gas or void per volume of solidmaterial in the matrix. The pore size of suitable tissue scaffoldsshould be at least about 50 microns, at least about 100 microns andpreferably at least about 150 microns, where pore size is roughly theaverage size of the largest dimension of the pores in the matrix.

A “hydrogel” is any porous matrix, including tissue scaffolds, where thepores are filled with a liquid. A hydrogel differs from a tissuescaffold in that a hydrogel may have a porosity of less than 60% and/ora pore size of less than 50 microns and cells typically do not growwithin such hydrogels of low porosity, although they may grow on thesurface of the same. While tissue scaffolds may be used as hydrogels incertain embodiments of the invention, in the typical embodiments, thehydrogel has a lower porosity and/or pore size as mentioned. Hydrogelsmay be made of the same scaffolding materials as tissue scaffolds, orother materials including but not limited to agars, agaroses, alginates,collagens, acrylamides, celluloses, starches, methacrylates and thelike.

“Scaffolding polymers” or “scaffolding materials” are the materials usedto make tissue scaffolds. The terms refer to both monomeric units of thematerials and the polymers made therefrom. Scaffolding polymers may bebiodegradable or non biodegradable. Suitable examples of biodegradableand non-biodegradable scaffolding polymers useful in the practice of theinvention are found in the description that follows, in the referencesincorporated herein, and in the references cited in the incorporatedreferences.

A “wafer” is a tissue scaffold formed into any suitable shape forimplantation into the tooth of a subject. Suitable wafer shapes include,but are not limited to microspheres, tubular barrels, concentric tubes,spiral tubes, cylindrical sponges, overlaying mats and the like.

A “hole of corresponding size” and “wafer of corresponding size” referto an indentation (hole) and a tissue scaffold wafer respectively, eachof a selected size, where the hole is formed in a tooth by ahole-forming process such as drilling, and the size of the hole is suchthat the scaffold wafer of selected size, when properly oriented, willfit into the hole with a movement tolerance in a lateral direction ofabout 0.01 to 0.5 mm and more preferably about 0.01 to about 0.1 mm.

Tooth Anatomy

A better understanding of the invention may be had with reference tobasic tooth anatomy as illustrated in FIG. 1. The tooth includes a crownregion 25 located above the gum 50, a neck region 35 in the vicinity ofthe gum 50, and a root region 45 located beneath the gum 50. Enamel 15covers the crown 25 of the tooth. Cementum 55 anchors the tooth in thebone 60 through the periodontal ligament 65. Dentin 20 is themicroscopically porous hard tissue under the enamel 15 and the cementum55 and is comprised of a combination of a porous mineral matrix andliving cells. Pulp tissue 70 is located beneath the dentin 20 in a pulpchamber that has a coronal region and a radicular (root canal) region 75ending at the foramen (hole) 90 at the end of the root where the pulptissue 70 becomes continuous with the periodontal ligament.

The dental pulp 70 consists of loose connective tissue derived fromectomesenchyrnal cells and is confined within the coronal region androot canals 75 of the tooth. The pulp 70 contains cells that provideodontogenic, nutritive, sensory, and defensive functions to the maturepulp 70 and allows for preservation of vitality during normalhomeostatic maintenance and during wound repair after injury.

The mature dental pulp 70 can be divided into two compartments: Theodontogenic zone and the pulp proper. The odontogenic zone includes theodontoblasts, which are the cells responsible for the production ofpredentin and dentin, the cell-free zone, the cell-rich zone, and theparietal plexus of nerves. The pulp 70 proper includes the majority ofthe remaining area of the pulp and consists primarily of fibroblasts andextracellular matrix, blood vessels, and nerves. As the pulp 70 ages,its volume decreases with a corresponding increase in dentin thickness.Numerous studies have demonstrated that the dental pulp 70 has aninherent capacity to respond to wounding in the absence of otherinflammatory insults.

The most predominant cell type in the dental pulp 70 is the fibroblast,but the pulp also contains odontoblasts, blood cells, Schwann cells,endothelial cells, pericytes and undifferentiated mesenchyma cells.Cells involved in the immune response, such as macrophages, mast cells,antigen processing cells (dendritic cells), and plasma cells can also befound in the pulp during periods of inflammation.

Odontoblasts are terminally differentiated, polarized pulpal cellsderived from the cranial neural crest, which are found in a peripherallayer closely associated with the predentin. The major function ofodontoblasts is production of unmineralized [predentin] and mineralizeddentin matrices.

The cell bodies form an irregularly columnar, epithelial-like layer onthe inner aspect of the dentin. The proximal surface of the cell body isadjacent to pulp cells, and the distal extremity is tapered and embeddedin predentin, an unmineralized layer of dentin-like material as well aswithin tubules in the mineralized dentin.

The major protein produced by the odontoblast is type I collagen and issecreted into the extracellular space at the predentin interface.Non-collagenous components of the extracellular matrix of predentin anddentin, including proteoglycans, glycosaminoglycans, phosphoproteins,and γ-carboxyglutamate-containing proteins, are also synthesized andsecreted by odontoblasts.

Functioning odontoblasts continue to produce predentin throughout thelife of the tooth. Odontoblasts retain the ability to upregulate proteinsynthetic activity in response to trauma after aging.

The remainder of the pulp 70 consists of a “stromal” tissue containingnerves, blood vessels, and lymphatics. The stromal tissue is composed ofcells and extracellular material. There appears to be only one type ofcell, resembling mesenchyme but capable of producing extracellularmaterial, including collagen. Thus, the cell can equally be termed afibroblast, or simply a pulpal cell.

Fibroblasts are the most numerous cells found in the dental pulp 70.They are stellate-shaped cells with long cytoplasmic extensions thatcontact adjacent fibroblasts or odontoblasts through gap-junctionalprocesses. Fibroblasts synthesize and secrete type I and type IIIcollagen, and other ECM components of the pulp, including proteoglycansand glycosaminoglycans. Collagen is the most abundant connective tissueprotein and occurs in several specific isotypes, types I through XII.Each is recognized as a specific genetic product differing in amino acidand polypeptide composition. In the pulp 70, type I and type III are themost abundant, with other types, such as IV and V being minorconstituents.

Fibroblasts or undifferentiated mesenchymal cells also have an importantrole in wound healing mechanisms in the pulp 70. The fibroblasts of thecell-rich zone are thought to differentiate into odontoblasts after theright stimulus—for example, growth factor, a bone morphogenic protein(BMP), a cytokine, or an inflammatory mediator, typically releasedduring wounding from the exposed predentin or dentin, or frominflammatory cells that have migrated to the wound site.

There are many other cells found in a vital dental pulp 70. Perivascular(a.k.a pericytes) cells are found in the dental pulp closely associatedwith the vasculature. These cells have been reported to be important inwound-healing mechanisms associated with pulpal repair mechanisms.Perivascular cells have also been shown to proliferate in response to aniatrogenic exposure of the dental pulp, and are thought to possiblyprovide replacement cells for the odontoblast layer in wounds where thecell-rich layer has been destroyed.

Endothelial cells line the lumen of the pulpal blood vessels andcontribute to the basal lamina by producing type IV collagen, anafibrillar collagen. They have been shown to proliferate after a pulpexposure in an attempt to neovascularize the wounded area during theprocess of wound healing.

Class II antigen processing cells have been demonstrated byimmunohistochemical methods in both the normal and inflamed pulp. Othervascular-derived cells found in the pulp during an inflammatorycondition include mast cells, B- and T-lymphocytes, polymorphonuclearneutrophils, and macrophages. These blood cells are of paramountimportance in fighting infection in the pulp because of the substancesthey contain: histamine, serotonin, cytokines, growth factors, and othercellular mediators. Schwann cells can also be found, which cellsenvelope nerve processes with a myelin sheath.

Dentin 20 forms the bulk of the tooth. It is lined on its outer aspectby enamel 15 on the crown 25 and by cementum 55 on the root 90. Theyoungest layer of dentin 20 formed at any particular time is adjacent tothe junction between the odontoblast cell body and its major process.This layer of young dentin 20 is essentially unmineralized, and endsabruptly in contact with the mature dentin. Mature dentin 20 isassociated with glycoproteins, an increased diameter of the collagenousfibrils, and a sudden and dramatic mineralization related to thesefibrils.

The outer surface of the root canal 75 is covered by a relatively thinlayer of the bone-like mineralized cementum 55. The cementum 55 is madeof a matrix of calcified collagenous fibrile, glycoproteins, andmucopolysaccharides. The outermost layer of cementum 55 is anuncalcified precementum produced by the discontinuous layer ofirregularly shaped cementoblasts.

The present invention is based on recognizing that one or more of theabove tissue types are involved in regenerating dentin 20 and thatcontacting viable pulp tissue 70 with a tissue scaffold in vivo,especially in combination with a morphogenic agent such as BMP-2, BMP-4or BMP-7, will initiate appropriate cell infusion into the tissuescaffold and ultimately result in cell proliferation as well asremineralization and regeneration of the dentin 20 and pulp 70.

Tissue Scaffolds Generally

The tissue scaffolds used in the present invention provide a matrix forthe cells to guide the process of tissue formation in vivo in threedimensions. Although the majority of mammalian cell types are anchoragedependent and will die if not provided an adhesion substrate, the tissuescaffolds of the present invention are not simply an adhesive substrate.Petri dishes and other non-tissue scaffold structures are generally usedin cell culture, whereby cell monolayers result, however, this does notlead to tissue regeneration. It will be understood that simple dishes,vials and other receptacles are not “tissue scaffolds” in the context ofthe present invention, although it is possible that the materials fromwhich they are formed may be adapted for this purpose by design and/orby combination with other tissue scaffold elements.

The materials utilized to fabricate a tissue scaffold for use in thepresent invention can generally be categorized into three types: (1)naturally derived materials, including ECM molecules, such as collagensand hyaluronic acid, (2) polysaccharides, such as alginate, agars,agaroses and the like that can be formed into hydrogels of sufficientporosity to serve as tissue scaffolds, and (3) bio compatible syntheticmaterials including any of a variety of polymers or co-polymers, whetherbiodegradable or non-biodegradable, that do not elicit adverse affectswhen implanted into tissue.

(1) Naturally-Derived Tissue Scaffolds.

Any one of a variety of naturally-derived tissue scaffold-like materialsmay be used to provide a framework for tissue growth in accordance withthe present invention. Because the tissue scaffold, or substantialportions thereof, are applied to a hole drilled in a subject's tooth,one will generally prefer to use a tissue scaffold that is derived froma biological tissue that is compatible with the tooth. Suchbiocompatibility requires that the tissue scaffold does not cause anysignificant adverse or untoward reactions when administered to thesubject. By using a biocompatible tissue scaffold significant immuneresponses and inflammatory reactions will be avoided.

A large number of naturally-derived tissue scaffold-like materials areavailable that may be used as tissue scaffolds in accordance with thisinvention, including those tissue scaffolds fabricated from human,animal or plant tissue. Potential advantages of these types of materialsare their biocompatibility and their biological activity. As many ofthese molecules are found within tissues, they may not induce anyforeign body reactions and are presumably receptive to the cell-mediatedremodeling that occurs during tissue repair and regeneration (Murphy etal., 1990; Yannas et al., 1989).

ECM molecules, such as collagen may be used as tissue scaffold materialsin certain practices of the invention. Type I collagen, the mostprevalent ECM molecule in the body, is readily isolated from animaltissues and has been extensively utilized to fabricate cell deliverydevices (Green et al., 1979; Yannas et al., 1981; Bell et al., 1981;Stem at al., 1990; Cavallaro et al., 1994). This material can beprocessed into a wide variety of structures for use in the invention,e.g., films, sponges and fibers (Green et al., 1979; Yannas et al.,1981; Bell at al., 1981; Stem et al., 1990; Cavallaro et al., 1994). Thestructure and resultant mechanical properties of collagen-basedscaffolds can be regulated by the process utilized to extract thecollagen from tissues (Cavallaro et al., 1994), and by variouscrosslinking processes. Collagen molecules may be crosslinked physicallyby dehydrothermal (Koide et al., 1993) or UV radiation treatments, orchemically by using various chemical agents (Cavallaro et al., 1994;Koide et al., 1993; DeLustro et al., 1990). However, the inflammatoryresponse to these materials and their erosion rate are dependent on thespecific cross-linking agent that is utilized (Cavallaro et al., 1994;Anselme, 1992; Koide et. al., 1993).

Suitable collagen tissue scaffolds are described, for example, in U.S.Pat. Nos. 4,347,234; 4,390,519; 4,394,370; 4,409,332; 4,538,603;4,585,797; 4,703,108; 4,837,285; 4,975,527; 5,081,106; 5,128,136;5,162,430; 5,197,977 and 5,206,028. Mineralized collagen, as disclosedin U.S. Pat. No. 5,231,169 may also be used in the present invention.

Type I collagen may also be combined with glycosaminoglycans to formgels which mimic native dermal tissue (Yannas et al., 1981; Stem et al.,1990; Heimbach et al., 1988). A variety of other ECM molecules,including laminin (Dixit, 1994; Guenard et al., 1992), have beenutilized as cell delivery tissue scaffolds, and any such tissue scaffoldmay be used in the context of the present invention.

(2) Polysaccharides

Polysaccharides may also be used as tissue scaffolds in accordance withthis invention. Alginate, a polysaccharide isolated from seaweed, haspreviously been used as a cell delivery vehicle. Water soluble sodiumalginate readily binds calcium, forming an insoluble calcium alginatehydrocolloid (Sutherland, 1991). These gentle gelling conditions havemade alginate a popular material to encapsulate cells fortransplantation (Lim and Sun, 1980; O'Shea at al., 1984; Ricordi et al.,1988; Sullivan et al., 1991; Lacy et al., 1991; Levesque et al., 1992;Soon-Shiong et al., 1994; Dixit, 1994; Kasai et al., 1994), and as aninjectable cell delivery vehicle (Atala et al., 1994).

The potential advantages of these natural materials have made thempopular for fabricating tissue engineering tissue scaffolds, and theymay certainly be used in the context of the present invention. However,these materials also have a number of disadvantages. Many of thesematerials are isolated from human or animal tissue, and are notavailable in large quantities. They suffer from large batch-to-batchvariations, and are typically expensive. Additionally, these materialsexhibit a limited range of physical properties (e.g., mechanicalstrength, erosion times). These drawbacks led the present inventors tocontemplate using synthetic materials to fabricate tissue scaffolds foruse in many aspects of this invention.

(3) Synthetic Tissue Scaffolds

Synthetic polymers are attractive scaffold materials as they can bereadily produced with a wide range of reproducible properties andstructures. Polymer tissue scaffolds also provide mechanical supportagainst compressive and tensile forces, thus maintaining the shape andintegrity of the scaffold in the environment of the tooth.

The morphology of the tissue scaffold can guide the structure of anengineered tissue (Vacanti et al., 1988), including the size, shape andvascularization of the tissue (Mooney et al., 1994a, 1994b, 1995b,1996a). The proper design of these tissue scaffolds allows them toexhibit the required range of mechanical and biological functions.Synthetic polymeric materials can be precisely controlled in materialproperties and quality. Moreover, synthetic polymers can be processedwith various techniques and supplied consistently in large quantities.The mechanical and physical properties of synthetic polymers can bereadily adjusted through variation of molecular structures so as tofulfill their functions without the use of either fillers or additives.Table 1 outlines different structural factors of polymers that can beused to adjust a variety of critical properties. TABLE 1 StructuralVariables used to Control Properties of Synthetic Biodegradable PolymersVariables Effects Examples Incorporation of May reduce/eliminateNon-immunologic PGA both natural and/ immunologic response and PLA (vs.collagens) or non-natural often found in naturally- monomers derivedpolymers Incorporation of Control kinetics of Hydrolyzable ester bondlabile groups in biodegradation in PGA polymer chain Incorporation ofControl chemical and Hydrophilic, hydrophobic functional groups physicalproperties of and amphiphilic in side chains polymers polyphosphazenesIncorporation of Control physical and Semi-crystalline L-PLA chiralcenters in mechanical properties of and amorphous D, L-PLA polymerchains polymer Possibility of Control properties of Glycolic and lacticacids utilizing multiple polymers PLGA monomers Use of naturalBiocompatible break- Lactic acid in PLA compounds as down productsmonomers Use of different Control physical and Branched polymers polymermechanical properties of exhibit lower viscosity architectures polymersthan linear onesAdapted from Wong and Mooney, 1997.

A variety of synthetic biodegradable polymers can be utilized tofabricate tissue scaffolds. In general, these materials are utilized asstructural elements in the scaffold. Poly(glycolic acid) (PGA),poly(lactic acid) (PLA), Poly lactic co-lactic acid (PLLA) andpoly(lactic acid)-poly(glycolic acid) (PLGA) polymers are commonly usedsynthetic polymers in tissue engineering. These polymers are alsoextensively utilized in other biomedical applications such as drugdelivery and are FDA approved for a variety of applications (Huang,1989).

A number of PGA, PLA, PLLA and PLGA and other synthetic polymer tissuescaffolds are known in the art, and are further described herein, anyone or more of which may be used in the context of the presentinvention. By way of example only, one may mention the PGA, PLA and PLGAformulations disclosed in any of: U.S. published applications No.2002/0119180A1, 2003/0031696 or U.S. Pat. Nos. 6,281,256, 6,472,210,5,885,829, 5,366,734; 5,366,733; 5,366,508; 5,360,610; 5,350,580;5,324,520; 5,324,519; 5,324,307; 5,320,624; 5,308,623; 5,288,496;5,281,419; 5,278,202; 5,278,201; 5,271,961; 5,268,178; 5,250,584;5,227,157; 5,192,741; 5,185,152; 5,171,217; 5,143,730; 5,133,755;5,108,755; 5,084,051; 5,080,665; 5,077,049; 5,051,272; 5,011,692;5,007,939; 5,004,602; 4,961,707; 4,938,763; 4,916,193; 4,898,734;4,898,186; 4,889,119; 4,844,854; 4,839,130; 4,818,542; 4,744,365;4,741,337; 4,623,588; 4,578,384; 4,568,559; 4,563,489; 4,539,981;4,530,449; 4,384,975; 4,300,565; 4,279,249; 4,243,775; 4,181,983;4,166,800; 4,137,921.

Where a tissue scaffold is to be administered to an oral tissue site,another reason for preferring a synthetic material is that the surfaceproperties of synthetic materials can be easily and reproduciblyaltered, as necessary. Plasma modification and grafting of relativelyinert substances, such as polyethylene oxide or polyvinyl alcohol, canmask the chemistry of the bulk tissue scaffold (Peppas and Langer,1994). The specific structure of adsorbed polymer coatings can becontrolled by varying the chemical structure and molecular weightpolydispersity of the coating polymer (Dan and Tirrell, 1993). Molecularself-assembly strategies can also be used to define the protein andcellular interactions with material surfaces (Prime and Whitesides,1991; Singhvi et al., 1994).

3a Biodegradable Tissue Scaffolds

Tissue scaffolds fabricated from biodegradable materials will erode overtime in the body to yield a completely natural tissue. These tissuescaffolds will not induce any chronic inflammatory responses, and cannotserve as a long-term site for infection. Biodegradable polymers havebeen utilized to engineer tissues ex vivo that will be structurallyintegrated with the host tissue (Langer and Vacanti, 1993; Heimbach etal., 1988; Hansbrough et al., 1992; Mooney et al., 1994a; 1994b; 1995a;1995b; 1996a; Johnson et al., 1994; Dixit, 1994; Kasai at al., 1994;Mooney and Vacanti, 1993). In addition, the use of synthetic,biodegradable tissue scaffolds will often be advantageous as thedegradation time of such synthetic tissue scaffolds can be designed tocoincide with the formation of a new tissue from the cultured cells.

While there are a variety of biodegradable polymers (Gilding, 1981;Peppas and Langer, 1994), polymers composed of monomers naturallypresent in the body (e.g., lactic acid, alpha-amino acids) are preferredfor use in certain aspects of the invention. These form aliphaticpolyesters of the poly(α-hydroxy acids) with the general formula->-O—CH(R)—CO-!- which derive from corresponding HO—CH(R)—COOH where R═Hin the case of glycolic acid (GA) and R═CH₃ in the case of lactic acid(LA), the latter being chiral, i.e., D- or L-isomer is possible. Thesepolymers have been used in bone osteosynthesis and reconstruction (Vertet al., 1984) and in drug delivery (Gombotz and Pettit, 1995).

Polymers of lactic acid, glycolic acid, and copolymers of the two havebeen utilized to fabricate tissue engineering tissue scaffolds (Heimbachet al., 1988; Hansbrough et al., 1992; Mooney et al., 1994a; 1994b;1995a; 1995b; 1996a; Johnson at al., 1994; Mooney and Vacanti, 1993).These polymers are readily processed into a variety of configurations,including fibers (Frazza and Schmitt, 1971), porous sponges (Mooney etal., 1995a; Mooney and Vacanti, 1993) and tubular structures (Mooney atal., 1995b).

The regular structure of homopolymers of lactic and glycolic acidresults in a crystalline structure (Gilding, 1981). Copolymerscontaining significant quantities of both monomers are amorphous(Gilding, 1981). This polymer family's widely varying mechanical anderosion properties (Table 2) results both from the varyingcrystallinity, and the differing hydrophobicity of lactic and glycolicacid (Gilding, 1981). Example tissue scaffolds with a variety ofdegradation times and mechanical properties suitable for the presentinvention are listed in Table 2. TABLE 2 Typical Yield Stress Values andErosion Times for Polymers of Lactic and Glycolic Acid YIELD TIME FOR50% STRESS EROSION POLYMER (Kpsi)* (weeks) polyglycolic acid 11.2 450/50 poly (D,L-lactic-co-glycolic 7.7 6 acid) 85/15 poly(D,L-lactic-co-glycolic 6.3 20 acid) poly (D,L-lactic acid) 6.6 35 Poly(L-lactic acid) 8.5 56Adapted from Wong and Mooney, 1997.*Values represent the mean of 5 measurements obtained using Instron.Data is adapted from Medisorb (Cincinnati, OH).product data.**The time at which ½ of the polymer has eroded(Polymer mass = ½ initialmass) following immersion in a buffered saline solution maintained at37.degree. C.

PGA and PLA can be prepared by two different routes, namely,polycondensation and ring opening polymerization (ROP). Generally, thesimple polycondensation is less expensive, but the resulting polymershave low and uncontrolled molecular weight and it is difficult toprepare copolymers (Gilding and Reed, 1979, Kricheldorf andKreiser-Saunders, 1996). It is believed that the antimony trioxidecatalyst typically used to effect polycondensation acts as bothpolymerizing and depolymerizing agent. Moreover, glycolic and lacticacids have a great tendency to cyclodimerize under these conditions andthis renders simple polycondensation an unsuitable method.

The preferred method for producing high molecular weight polymers is ROPof the cyclic dimer, glycolide (and/or lactide). Depending on thecatalyst involved, three different mechanisms have been reported:cationic, anionic and insertion. Among these, the insertion mechanismusing metal alkoxides or carboxylates is the most desirable pathway andis the choice in commercial production (Frazza and Schmitt, 1971).Typical examples of catalysts of this class are aluminum, zinc,titanium, zirconium, antimony, tin(IV) and tin(II) alkoxides orcarboxylates.

The insertion mechanism allows the preparation of high molecular weightpolylactides without racemization up to temperatures above 150° C.(Kricheldorf and Kreiser-Saunders, 1996). The mechanism of ROP oflactones has been reviewed by Penczek and Slomkowski (1989). The tincatalyst, tin(II) octoate, was used extensively because of itsacceptance by the FDA as a food stabilizer.

The polymerization of glycolide can be carried out in bulk at 220° C.for 4 h, at which time a 96% conversion and molecular weights from 10⁴to 10⁶ have been reported. Copolymerization of glycolide with lactidehas also been investigated. The reactivity ratios at 200° C. have beenfound to be 2.8 for glycolide and 0.2 for lactide. This indicates thatcopolymers of glycolic and lactic acids will have broad compositionalranges, with glycolide always being preferentially polymerized at lowconversions and lactide being incorporated to ever-increasing extents asthe glycolide is depleted (Gilding and Reed, 1979).

During the advanced stages of most ROPs, additional reactions such asester-ester interchange and chain unzipping may take place. The extentsof these reactions are affected by the reactivity of the ester moieties.These events can have a significant effect on the composition of thefinal product. Due to the ester exchange, cyclic dimer, trimer and, to aless extent, cyclic oligomers could be found along with the reformedmonomer. In the case of copolymerization, additional randomization ofthe polymer chain would occur as a consequence of the ester-esterexchange of different ester moieties (Shalaby and Johnson, 1994). Themicrostructures of PLGA copolymers can be determined by both proton andcarbon NMR spectroscopy (Ksaperczyk, 1996). It has been reported thatthe block lengths increase and, at the same time, the extent oftransesterification decreases with decreasing polymerizationtemperature.

PGA was first developed as the synthetic absorbable suture, Dexon(Frazza and Schmitt, 1971). PGA has high crystallinilty, a high meltingpoint and low solubility in organic solvents (Table 3). The polymer(fiber grade, inherent viscosity=1.2-1.6 dL/g in hexafluoroisopropanol)can be spun into multifilament or monofilament yarns for the productionof braided and monofilament sutures, respectively (Frazza and Schmitt,1971; Chujo et al., 1967). A typical suture braid has a tensile strengthof 80-100 Kpsi (Table 4). Owing to the hydrophilic nature of PGA, Dexonsutures tend to lose their mechanical strength rapidly (50%) over aperiod of 2 weeks and are absorbed in about 4 weeks after implantation(Frazza and Schmitt, 1971; Reed and Gilding, 1981; Katz and Turner,1970). TABLE 3 Crystallinity and Thermal Properties of PGA, PLA andCopolymers Polymer % Crystallinity Tm Tg PGA 46-52 225 36 90:10 PGLA 40210 37 50:50 PGLA 0 None 55 PLA 37 185 57 dl-PLA 0 None NIAAdapted from Wong and Mooney, 1997.

TABLE 4 Mechanical Properties of PGA and 90:10 PGLA 90:10 PGLA PGA(Dexon) (Vicryl) Tensile strength (Kpsi) 106 95 Knot strength (Kpsi) 6563 Elongation (%) 24 25Adapted from Wong and Mooney, 1997.

The presence of an extra methyl group poly(L-lactic acid) (PLA) orpoly(D-lactic acid) (d-PLA) makes them more hydrophobic than PGA. Forinstance, films of PLA only take up approximately 2% water (Gilding andReed, 1979). In addition, the ester bond in PLA is less labile tohydrolysis due to steric hindrance of the methyl group. Therefore, PLAdegrades much slower than PGA (Reed and Gilding, 1981) and has highersolubility in organic solvents.

PLA is employed much more often than d-PLA, since the hydrolysis of PLAyields L-lactic acid which is the naturally occurring stereoisomer oflactic acid. Whereas PLA possesses about 37% crystallinity, theoptically inactive poly(DL-lactic acid) (dl-PLA) is amorphous. Thedifference in the crystallinity of dl-PLA and PLA has importantpractical outgrowths. For instance, the amorphous dl-PLA is usuallyconsidered in drug delivery application where a homogeneous dispersionof the active species within a monophasic matrix is desired (Engelbergand Kohn, 1991). However, the semicrystalline PLA is preferred in caseswhere high mechanical strength and toughness are required, for example,in orthopedic devices (Leenslag at al., 1987, Vainionpaa et al., 1987;Hay et al., 1988). It is pertinent to note that γ-irradiation of PLAcauses chain scission, crosslinking and a decrease in crystallinity(Gupta and Deshmuth, 1983). Therefore, caution should be taken whensterilizing the polymer matrices by γ-irradiation.

To widen the range of materials properties exhibited by PGA, copolymersof GA and LA (PLGA) have been studied. Whereas PGA is highlycrystalline, PLGA usually exhibit lower crystallinity and Tm (Gildingand Reed, 1979). For example, while PGA and PLA are partiallycrystalline, 50:50 PLGA is entirely amorphous. These morphologicalchanges result in an increase in the rates of hydration and hydrolysis.Thus, copolymers tend to degrade more rapidly than PGA and PLA (Mooneyat al., 1995b).

The degradation mechanism of PGA and copolymers in vitro is usuallyregarded as bulk erosion (Gombotz and Pettit, 1995). This is evidentfrom the fact that a significant molecular weight decrease usuallyprecedes monomer release from the polymer samples. This mechanism ofdegradation may be undesirable in certain applications. The relativelyrapid release of large quantities of acid (glycolic and/or lactic acids)may lead to a local acidosis if a large mass of these polymers ispresent in a concentrated form (e.g., a solid pin). However, highlyporous scaffolds are utilized in the present invention are highly porousand therefore contain a relatively low mass of polymer per unit volume.The highly porous structure of the scaffolds assists cell penetration aswell as polymer degradation (Mooney at al., 1994b; 1995a; 1995b). Therate of degradation will be affected by the morphology of the scaffoldand the large surface areas speed up the diffusion of water moleculesinto the bulk of the polymers when they are placed in an aqueousenvironment (e.g., in vivo).

The polymers undergo random chain scission by simple hydrolysis of theester bond linkage and the monomer diffuses out of the polymer bulk intowater (Reed and Gilding, 1981). It is important to note that loss ofmechanical strength of PGA is faster when the polymer is incubated at atemperature higher than its Tg. This indicates that the glassy stateprotects PGA from hydrolysis since all short term chain motions arefrozen. Water diffusion, and therefore hydrolysis, is more facile attemperatures above Tg. It is also relevant to mention that the Tg's ofPGA and some copolymers are very close to the physiological temperature.Polymeric materials may undergo significant structural change afterimplantation due to water penetration and loss of mechanical strength.It is also speculated that enzymatic action may partially contribute tobiodegradation of PGA in vivo.

The chemical compositions and the ratio of monomers used in thepolymerization reaction strongly influence the degradationcharacteristics of the copolymer. The degradation rates for copolymersof GA and LA have been shown to be influenced by factors that affectpolymer chain packing, i.e., crystallinity, and hydrophobicity. Sincedegradation is induced by hydrolysis, a crystalline structure orhydrophobic polymer composition disfavors dissolution and degradation.

Gombotz and Pettit (1995) summarized the specific factors affectingcopolymer crystallinity and hydrophobicity: (i) the ratio of lactide toglycolide monomer in the copolymer, (ii) the stereoregularity of themonomer units in the polymer affects polymer chain packing, (iii)randomness of lactide and glycolide decrease the ability of chains tocrystallize, and (iv) low molecular weight polymers degrade faster thanhigh molecular weight polymers, especially when the end groups are freeacid rather than capped with ester or other groups. Mass loss frompolymer samples comprised of PLA is insignificant in the experimentaltime period (about 50 weeks). However, those comprised of copolymers ofGA and LA or dl-PLA degrade much faster (the higher the glycolic acidcontent, the higher the degradation rate) (Mooney et al., 1995b).

The presence of monomers and low molecular weight cyclic oligomers inabsorbable polymers should be avoided, for they degrade much morerapidly than the polymers and can lead to undesirable chemical andbiological effects. (Shalaby and Johnson, 1994) It has been shown thatpolylactide with increased monomer content exhibits a higher rate ofbioabsorption and a more drastic decrease of molecular weight. (Nakamuraet al., 1989)

Those of skill in the art will understand that the PLA, PGA and PLGApolymers are just one example of biodegradable polymer tissue scaffoldsthat may be used in this invention. Further biodegradable tissuescaffolds include polyanhydrides, polyorthoesters, and poly(amino acids)(Peppas and Langer, 1994). Any such polymer materials may be utilized tofabricate a biodegradable polymer tissue scaffold with controlledproperties for use in this invention. Further biodegradable polymersthat produce non-toxic degradation products are listed in Table 5. TABLE5 Example Polymers Recognized as Biodegradable Synthetic PolypeptidesPolydepsipeptides Nylon-2/nylon-6 copolyamides Aliphatic polyestersPoly(glycolic acid)(PGA) and copolymers Poly(lactic acid)(PLA) andcopolymer Poly(alkylene succinates) Poly(hydroxy butyrate) (PHB)Poly(butylene diglycolate) Poly(.epsilon.-caprolactone) and copolymersPolydihydropyrans Polyphosphazenes Poly(ortho ester) Poly(cyanoacrylates) Natural Modified polysaccharides cellulose, starch, chitinModified proteins collagen, fibrinAdapted from Wong and Mooney, 1997.

Little modification of these polymers is possible because there are noother functional groups on the side chain, except the methyl of thelactic acid residue. One possibility to modify the properties of thesepolymers is to form copolymers with residues having more diverse sidechain structures, e.g., lysine.

A new monomer,3-(Ne-benzoxycarbonyl-L-lysyl)-6-L-methyl-2,5-morpholinedione, was bulkcopolymerized with L,L-lactide in the presence of stannous octoate ascatalyst using the same ROP techniques utilized for lactide andglycolide (Barrera et al., 1993). The lysine content was determined byNMR technique to be approximately 1.3 mole %.

A poly(lactide-co-lysine) functionalized with peptide containing thearginine-glycine-aspartate (RGD) sequence was prepared by removal of thebenzyoxycarbonyl protecting group on the lysyl residue and peptidecoupling. The peptide concentration was found to be approximately 3.1mmol/g which could be translated into a peptide surface density of 310fmol/cm². A surface density of as low as 1 fmol/cm² of an RGD peptidehas been previously determined to promote cell adhesion to an otherwisenonadherent surface (Massia and Hubbell, 1991). Therefore, by carefullyprocessing the copolymer, biodegradable films with cell adheringproperties can be prepared from the copolymer of lactide and lysine.

Other strategies have also been employed to widen the properties ofpolylactides. For example, PLA has also been synthesized as an acrylicmacromonomer and subsequently copolymerized with polar acrylic monomers(e.g., 2-hydroxyethylmethacrylate) (Barakat et al., 1996). Thesepolymers were studied as amphiphilic graft copolymers for drug deliverypurposes. The surface properties of these polymers may be controlled bythe ratio of the PLA graft length and copolymer content, and can bepotentially used to control the drug release profile andbiodistribution. Other examples of this approach include grafting PLAblocks to geraniol and pregnenolone (Kricheldorf and Kreiser-Saunders,1996).

Other Polyesters

Properties of polyesters can also be varied by changing the structuresof the polymer backbones. Polycaprolactone (PCL), having two more carbonatoms than PGA on the polymer backbone, has been studied as a substratefor biodegradation and as a matrix for drug release systems (Huang,1989). Its degradation in vivo is much slower than PGA, therefore, it issuitable for controlled release devices with long in vivo life times(1-2 years). PCL can be prepared by anionic ring opening polymerizationof ε-caprolactone using metal hydroxide initiators (Jerome and Teyssie,1989).

Poly-β-hydroxy acid can be prepared by both cationic and anionic ringopening polymerizations (Penczek and Slomkowski, 1989; Jerome andTeyssie, 1989). For example, 100% syndiotacticpoly(β-DL-hydroxybutyrate) has been prepared by treating thecorresponding lactone with cyclic dibutyltin initiators to yield highmolecular weight polymers (Kricheldorf and Lee, 1995). Bacteria alsoproduce the chiral, isotactic poly(β-D-hydroxybutyrate) as a highlycrystalline biopolymer (Holmes, 1988).

Numerous analogs of poly(β-hydroxy acid) have been synthesized eitherchemically by the ring opening polymerization or biologically by feedingunusual carbon sources to bacteria (Timmins and Lenz, 1994). Themicrobial synthesis of polyesters has been reviewed by Gross (1994). Dueto their biocompatibility and biodegradability, different blends ofpolycaprolactone, poly (β-hydroxybutyrate) and other polymers have beenfabricated for medical devices (Yasin and Tighe 1992), drug deliveryapplications (Wang, 1989) and cell microencapsulation (Giunchedi et al.,1994; Embleton and Tighe, 1993).

Surface-eroding polymer matrices are attractive for a variety of tissueengineering applications. The monomer release would be steady over thelifetime of the matrices in contrast to PLA and PGA. In addition, thegradual loss of polymer from the surface of the scaffold may allow thesurrounding tissue to serially fill the space vacated by the polymer.

Polyorthoesters are an example of surface-eroding polymers. Thehydrophobic character of the polymer limits water penetration andhydrolysis to only the exterior surface of the polymer matrix (Heller,1985). Thus the surface erosion is much faster than that of the bulk.The chemical and physical properties of polyorthoesters have beenreviewed (Heller and Daniels, 1994) and depend on the chemicalstructures of the constituent monomers. For example, reaction ofbis(ketene acetal) with rigid trans-cyclohexane dimethanol produce arigid polymer with a Tg of 10° C., whereas that of the flexible diol1,6-hexanediol produces a soft material having a Tg of 20° C. Mixture ofthe two diol results in polymers having intermediate Tg.

Degradation of the polymers is acid-induced, and degradation rates canbe increased by adding acidic excipients or by increasing thehydrophilicity of the polymer matrix. Conversely, degradation can beretarded by using basic excipients such as Mg(OH)₂ (Gombotz and Pettit,1995). Current applications of these polymers include sustained drugdelivery as well as hard and soft tissue fixation (Helier and Daniels,1994).

Polyorthoformate, polycarbonate, poly(oxyethylene glycolate),poly(1,4-butylene diglycolate) and polyurethane are other biodegradablepolymers that may have applications in tissue engineering. Many of thesepolymers have been previously utilized as drug delivery matrices (Huang,1989).

There are therefore a large number of polyesters and analogs that arebiodegradable. Their mechanical properties can be controlled largely bythe chemical structures of the constituent building blocks and can bevaried from tough to elastic. The biocompatibility of these polymers ispresumed to result from non-toxic degradation products. Bioactiveelements can be attached to this class of materials in order to mimicnatural extracellular matrix molecules.

Polypeptides

Proteins, one of the most important biomolecules in nature, belong tothis class of biopolymers. However, polypeptides of a single amino acidor copolymers were generally regarded as impractical industrialmaterials (Nathan and Kohn, 1994). Amino acid N-carboxyanhydrides wereprepared as the monomeric starting materials, and this addedconsiderably to the cost of all polypeptides. These polymers were thusexpensive even if they were derived from cheap amino acids. In addition,it was almost impossible to control the sequence of the protein polymersusing random copolymerization techniques. Most polypeptides areinsoluble in common organic solvents. The need for exotic solventsystems to process these materials combined with their thermalinstability made them poor engineering materials.

A number of recent approaches may, however, bypass these difficulties.Advances in genetic engineering have enabled investigators to obtainprotein polymers by inserting DNA templates of predetermined sequencesinto the genome of bacteria. Collagen-like, silk-like, andsilk-elastin-like proteins have been synthesized by this technique(Goldberg et al., 1989; Cappello et al., 1990; McGrath et al., 1992).

The general concept (O'Brien, 1993) involves the incorporation of aminoacid sequences with desired properties, e.g., cell adhesion orelasticity, into the protein polymers to produce materials ofpredetermined structure and controlled properties. For example, a celladhering sequence of RGD has been incorporated into silk-like proteinpolymers in a manner such that the tripeptide sequence is exposed forcell attachment (Tirrell et al., 1994).

Investigators have also developed chemical synthetic techniques that arecomplementary to the genetic approach to prepare such materials. Forinstance, different rigid non-peptide, organic segments have beencombined with leucine-glutamine-proline, a sequence of the calciumbinding domain of bovine amelogenins, using a completely syntheticapproach (Sogah et al., 1994). The advantage of this class ofprotein-based hybrid polymers is the virtually unlimited choice ofbuilding blocks for the polymers. In contrast, genetically engineeredproteins can only make use of the 20 natural and a limited number ofunnatural amino acids for the construction of polymers.

Rigid organic segments have been used to reduce the conformationalflexibility of 30 the peptide chain through the formation of peptidesecondary structures (e.g., β-sheet or β-turns). Controlled folding ofthe polymer backbone has been reported using such ordered buildingblocks (Wong, 1996). The potential application of these materials totissue engineering is significant. These synthetic techniques allowprecise control of material properties, while maintaining the freedomand flexibility to design protein-like materials with desirablebiological and chemical properties. These properties may make this classof materials desirable matrices for tissue engineering applications.

Urry and coworkers have also studied elastin protein-based polymers asbiocompatible materials. These polymers, also known as bioelasticmaterials, are elastomeric polypeptides comprised of the generalrepeating sequences glycine-any amino acid-glycine-valine-proline(GXGVP). The polymers were synthesized by the self condensation of theactivated p-nitrophenol ester of the pentapeptide building blocks(Prasad et al., 1985). The molecular weight of these polymers areconsidered to be higher than 50,000. The polymers can be cross-linked byγ-irradiation to form an insoluble matrix without detectable residuedestruction. Cell adhesion sequences (e.g., RGD) and enzymatic siteshave been incorporated into the polymers for cell attachment andcatalytic activity studies, respectively. Synthesis utilizing geneticengineering approach has also been reported. These class of polymershave been reported to exhibit excellent biocompatibility (Urry et al.,1995; Urry, 1993).

One specific polypeptide (GVGVP)N has been shown to undergo an inversetemperature transition in water (Urry, 1988a; Urry, 1988b). Themechanism of such elasticity has been demonstrated to be entropic innature and is apparently due to the internal chain dynamics of theordered polypeptide structure. This is contrary to the common beliefthat the elasticity of elastin, similar to synthetic polymers, is due torandom chain network and random end-to-end distances (Alberts et al.,1983). The transition temperature can be controlled by the amino acidcomposition, pH and phosphorylation, electrochemical, photochemical andchemical reactions of prosthetic groups. Therefore, a device thatconverts chemical energy into mechanical work can be constructed.

Non-Biodegradable Tissue Scaffolds

Although biodegradable tissue scaffolds will have advantages in certainembodiments, they are by no means required for use in practicing thepresent invention. Non biodegradable tissue scaffolds are suitable ifthey are biocompatible and can promote deposition of ECM molecules andbe mineralized in vivo.

Calcium phosphate ceramics are non-biodegradable tissue scaffolds thatare extensively used in engineering bone tissue (Ducheyne, 1988) and maybe used in the present invention. A suitable ceramic that may be used isdescribed, for example, in U.S. Pat. No. 4,596,574.

Both hydroxyapatite and tricalcium phosphate, and mixtures of the two,may be utilized. These materials only release calcium and phosphate asbreakdown products. They display no local or systemic toxicity, andbecome directly bonded to adjacent bone tissue with no interveningfibrous capsule (Ducheyne, 1988). The erosion and mechanical propertiesof these materials are controlled by the specific chemical compositionand processing conditions (Lemons, 1988).

Applications of hydroxyapatite, tricalcium phosphate, and mixturesthereof are currently limited by their brittle nature and generally poormechanical properties (Jarcho, 1981). However, this is not a drawback inthe context of the present invention. In various aspects, calciumphosphate and other minerals such as fluoride, are intentionallyassociated with the tissue scaffolds. Such pre-mineralization promotesproper regeneration and remineralization of dentin.

Other non-biodegradable polymers include semipermeable polymers such aspoly(acrylonitrile-co-vinyl chloride) (Emerich et al., 1992; Sagan etal., 1993; Guenard et al., 1992), polylysine (Lim and Sun, 1980; O'Sheaet al., 1984; Ricordi et al., 1988; Sullivan et al., 1991; Lacy et al.,1991; Levesque et al., 1992; Soon-Shiong at al., 1994), celluloseacetate (Yang et al., 1994) and polysulfone (Yang at al., 1994).Although generally intended for use in immobilized cells, the use ofsuch polymers in the context of the present invention is certainly notexcluded. These polymers may also be used with a variety of gels,including alginate and polyphosphazenes.

Hydrogels

As noted in foregoing definitions above, a hydrogel and a tissuescaffold have interchangeable meanings if the hydrogel is formed withsufficient porosity and pore size to allow three dimensional cell growthwithin the matrix of the hydrogel. Accordingly, the hydrogels describedherein are also suitable for making tissue scaffolds when the hydrogelsare formed with sufficient porosity and pore size to permit growth oftissue within the hydrogel matrix. However, some hydrogels have lessporosity than required of a tissue scaffold and are used as a sealingbarrier to cover implanted tissue scaffolds in certain aspects of theinvention.

Blends, Interpenetrating Networks (IPN) and Composites

The use of polymer blends or composites (polymeric composite materials)as biomaterials is a concept that nature exploits in assembling the ECMin tissue. The ECM of tissues typically contains a composite ofdifferent macromolecules and non-macromolecular materials. For example,glycoaminoglycans, which are usually covalently linked to proteins toform proteoglycans, constitutes a gel-like, highly hydrated structuresubstance in the which the collagen fibers are embedded (Wight et al.,1991; Giusti et al., 1993).

Blends of fibrin and polyurethane have previously been formed by acombined phase-inversion and spray process to produce highly poroussmall-diameter vascular prostheses. (Giusti et al., 1985, Soldani etal., 1992). These materials exhibit high thermal stability, and theirtensile behavior ranged from that of an elastic polyurethane tube tothat of a natural blood vessel. Hydrogels of fibrin and poly(vinylalcohol), blends or IPNs of collagen and poly(vinyl alcohol), blends ofhyaluronic acid with poly(vinyl alcohol) or poly(acrylic acid), andblends based on esters of hyaluronic acid have been reported (Giusti etal., 1993). These materials may be suitable for a variety ofapplications including soft tissue replacement, drug delivery,nerve-guide growth and cardiovascular devices. This class of materialshas great potential owing to the large number of readily availablesynthetic polymers that can be mixed with biopolymers.

Polysaccharides

Polysaccharides are carbohydrates characterized by the presence of arepeating structure in which the interunit linkages are of theO-glycoside type. The hydrophilicity of polysaccharides, along with theease in which they can be formed into hydrogels, makes these materialsideal for many tissue engineering applications in which one desires toimmobilize cells within a matrix. The variety of saccharides monomers(about 200) and the variety of possible O-glycoside linkages result in adiversity of polysaccharide structures and conformations.Polysaccharides may be derived from different sources including plants(starch, cellulose), animal (glycogen), algae and seaweeds (alginate andagarose) and microorganisms. These materials are usually considered asnaturally-derived products. However, since polysaccharides are widelyutilized as immobilization materials, they are used here as standards towhich other synthetic materials are compared.

a. Algal Polysaccharides: Alginate and Agarose

Algal polysaccharides have been the most commonly utilized hydrogelmaterials. This is due to their gentle gelling conditions, widespreadavailability, and relative biocompatibility. The main starting sourcesof alginate are species of brown algae (Phaeophyceae). The algae aretypically subjected to a number of processing steps to produce purealginate which is the major polysaccharide present and may comprise upto 40% of the dry weight. It is part of the intracellular matrix andexists, in the native state, as a mixed salt of the various cationsfound in sea water (e.g., Mg²+, Ca²⁺, Sr²⁺, Ba²⁺ and Na⁺). Due toselectivity of cation binding, the native alginate is mainly found inthe insoluble gel form, which results from cross-linking of alginatechains by Ca²⁺.

All alginates are copolymers of D-mannuronate (M) and L-guluronate (G).However, alginates from different algal sources have differentcompositions, and thus, different physical and mechanical properties.The block length of monomer units, and overall composition of thealginate and molecular weight, determine the properties of alginates.For example, calcium alginates rich in G are stiff materials(Sutherland, 1991).

Alginate selectively binds divalent metal ions such as Ba²⁺, Sr²⁺ andCa²⁺. The binding selectivity increases with G content, andpolymannuronate is essentially non-selective. The calcium ions are,therefore, selectively bound between sequences of polyguluronateresidue, and are held between diaxially linked L-guluronate residueswhich are in the ¹C₄ chair conformation. The calcium ions are thuspacked into the interstices between polyguluronate chains associatedpairwise and this structure is named the “egg-box” sequence. The abilityto form a junction zone depends on the length of the G-blocks indifferent alginates. Since the mechanical strength of alginate gelsdepend on the block lengths and M/G content, there have been efforts tomodify the M/G ratio by alginase to increase the G content (Skjak-Braeket al., 1986). It is expected that chemically modified alginate wouldalso produce materials of desirable properties. For example, bacterialalginates that contains acetyl groups generally exhibit differentphysical and mechanical properties from those of algal sources (Ott andDay, 1995).

Alginate can be gelled under mild conditions, allowing cellimmobilization with little damage. Binding of Mg²⁺ and monovalent ionsto alginate does not induce gelation of alginate in aqueous solution(Sutherland, 1991). However, exposure of alginate to soluble calciumleads to a preferential binding of calcium and subsequent gelling. Thesegentle gelling conditions are in contrast to the large temperature orsolvent changes typically required to induce similar phase changes inmost materials.

Alginates have been utilized as immobilization matrices for cell(Smidsrod and Skjak-Braek, 1990), as an injectable matrix forengineering cartilaginous tissue to treat vesicoureteral reflux invarious animal models (Atala et al., 1993 and Atala et al, 1994), and asinjectable microcapsules containing islet cells to treat animal modelsof diabetes (Sun et al., 1984).

The open lattice structure and wide distribution of pore sizes incalcium alginate preclude the controlled release of large molecules(e.g., proteins) from these materials and limits the use of purealginate for entrapment of whole cells or cell organelles (Smidsrod andSkjak-Braek, 1990). However, alginate membrane can be modified byincorporating other polymeric elements (e.g., lysine, poly(ethyleneglycol), poly(vinyl alcohol) or chitosan) (Polk et al., 1994, Kung etal., 1995). These modified systems have been used to control the releaseof proteins from alginate beads. Haemostatic swabs made of calciumalginate have also been clinically utilized to reduce blood loss duringsurgical procedures. The calcium ions in alginate may assist the bloodclotting process by activating platelets and clotting factor VII (Blairet al., 1990).

Agarose is another type of marine algal polysaccharide. In contrast toalginate, agarose forms thermally reversible gels. Agarose will set atconcentrations in excess of 0.1%, depending on the sulfate content, andat temperatures considerably below (about 40° C.) the gel-meltingtemperature (about 90° C.). The latter parameter is correlated to themethoxy content. The proposed gel structure is bundles of associateddouble helices and the junction zones consist of multiple chainaggregations (Yalpani, 1988). Agarose has been used largely in gels forelectrophoresis of proteins and nucleic acids. However, agarose gelshave also been used as supporting materials for electrophoresis ofbacteriophages (Serwer, 1987) and migration studies of leukocytes(Kallen et al., 1977). Although applications in tissue engineering havenot been reported, its adjustable gelling behaviour may render lowtemperature melting agarose a suitable injectable and immobilizationmatrix material.

b. Other Polysaccharides

Microbial polysaccharides are ubiquitous in nature and very abundantbiopolymers. They are of interest because of their unusual and usefulfunctional properties. Some of these properties are summarized by Kaplanet al. (1994) as: (i) film-forming and gel-forming capabilities, (ii)stability over broad temperature ranges, (iii) biocompatibility (naturalproducts avoid the release/leaching of toxic metals, residual chemicals,catalyst, or additives), (iv) unusual rheological properties, (v)biodegradability, (vi) water solubility in the native state or reducedsolubility if chemically modified, and (vii) thermal processability forsome of these polymers. Some examples of microbial polysaccharidessuitable for forming hydrogels for use in this invention are listed inTable 6. It is worthy to note that gellan, one of the microbialpolysaccharides, has been investigated as an immobilization material forenzymes and cells (Doner and Douds, 1995). TABLE 6 Some PolysaccharidesSynthesized by Microorganisms Polymer Structure Fungal Pullulan (N)1,4-1,6-α-D-Glucan Scleroglucan (N) 1,3-1,6-αD-Glucan Chitin (N)1,4-β-D-Acetyl glucosamine Chitosan (C) 1,4-β-D-N-Glucosamine Elsinan(N) 1,4-1,3-α-D-Glucan Bacterial Xanthan gum (A) 1,4-β-D-Glucan withD-mannose; D- glucuronic acid as side groups Curdlan (N) 1,3-β-D-Glucan(with branching) Dextran (N) 1,6-α-D-Glucan with some 1,2-1,3-1,4-α-linkages Gellan (A) 1,4-β-D-Glucan with rhamose, D- glucuronicacid Levan (N) 2,6-β-D-Fructan with some β-2,1- branching Emuisan (A)Lipoheteropolysaccharide Cellulose (N) 1,4-β-D-Glucana N = neutral, A = anionic and C = cationic.Adapted from Wong and Mooney, 1997.

Non-Natural Hydrogels

a. Polyphosphazenes

Polyphosphazenes contain inorganic backbones comprised of alternatingsingle and double bonds between nitrogen and phosphorus atoms, incontrast to the carbon-carbon backbone in most other polymers. Theuniqueness of polyphosphazenes stems from the combination of thisinorganic backbone with versatile side chain functionalities that can betailored for different applications. The degradation of polyphosphazenesresults in the release of phosphate and ammonium ions along with theside groups (Allcock, 1989; Scopelianos, 1994).

Linear, uncross-linked polymers can be prepared by thermal ring openingpolymerization of (NPCl₂)₃ and the chloro group replaced by amines,alkoxides or organometallic reagents to form hydrolytically stable, highmolecular weight poly(organophosphazenes). Depending on the propertiesof the side groups, the polyphosphazenes can be hydrophobic, hydrophilicor amphiphilic. The polymers can be fabricated into films, membranes andhydrogels for biomedical applications by cross-linking or grafting (Loraet al., 1991; Allcock et al., 1988; Allcock, 1989). Bioerodible polymersfor drug delivery devices have been prepared by incorporating hydrolyticside chains of imidazole (Laurencin et al., 1987) for skeletal tissueregeneration (Laurencin at al., 1993). Non-degradable phosphazenes havebeen used as denture liner (Razavi et al., 1993). Their use in thepresent invention is thus particularly contemplated.

b. Poly(Vinyl Alcohol) (PVA)

PVA is not synthesized directly but is the deacetylated product ofpoly(vinyl acetate). Polyvinyl acetate is usually prepared by radicalpolymerization of vinyl acetate (bulk, solution or emulsionpolymerizations) (Finch, 1973). PVA is formed by either alcoholysis,hydrolysis or aminolysis processes of poly(vinyl acetate). Thehydrophilicity and water solubility of PVA can be readily controlled bythe extent of hydrolysis and molecular weight. PVA has been widely usedas thickening and wetting agent.

PVA gels can be prepared by cross-linking with formaldehyde in thepresence of sulfuric acid (Schwartz et al., 1960). Theseformaldehyde-cross-linked PVA materials have been used as prosthesis fora variety of plastic surgery applications including breast augmentation(Clarkson, 1960 and Peters and Smith, 1981), diaphragm replacement(Haupt and Myers, 1960) and bone replacement (Camerson and Lawson,1980). However, a variety of complications were found after long termimplantation, including calcification of the PVA (Peters and Smith,1981). In the present invention, calcification is a desired featurebecause when a tissue scaffold is implanted in the tooth of a subject,regeneration of the dentin or enamel involves remineralization of tissuescaffold with calcium phosphate.

More recently, PVA was made into an insoluble gel using a physicalcross-linking process. These gels were prepared with a repeatedfreezing-thawing process. This causes structural densification of thehydrogel due to the formation of semicrystalline structures. The use ofthis gel in drug delivery applications has been reported (Peppas andScott, 1992; Ficek and Peppas, 1993). However, PVA is not trulybiodegradable due to the lack of labile bonds within the polymer bond.Only low molecular weight materials are advisable to be used as implantmaterials.

c. Poly(Ethylene Oxide) (PEO)

PEO or polyethylene glycol can be produced by the anionic or cationicpolymerization of ethylene oxide using a variety of initiators (Boileau,1989; Penczek and Kubisa, 1989). PEO is highly hydrophilic andbiocompatible, and has been utilized in a variety of biomedicalapplications including preparation of biologically relevant conjugates(Zalipsky, 1995), induction of cell membrane fusion (Lentz, 1994) andsurface modification of biomaterials (Amiji and Park, 1993). Differentpolymer architectures have been synthesized and some of theirapplications in medicine have been recently reviewed (Merrill, 1993).For example, PEO can be made into hydrogels by γ-ray or electron beamirradiation and chemical crosslinking (Cima et al., 1995; Belcheva etal., 1996). These hydrogels have been used as matrices for drug deliveryand cell adhesion studies.

d. Pluronics

Pluronic polyols or polyoxamers are block copolymers of PEO andpoly(propylene oxide and are usually synthesized by anionicpolymerization in the form of a ABA triblock using a difunctionalinitiator (Schmolka, 1972). Pluronics F 127, which contains 70% ethyleneoxide and 30% propylene oxide by weight with an average molecular weightof 11,500, is the most commonly used gel-forming polymer matrix todeliver proteins (Gombotz and Pettit, 1995).

This polymer exhibits a reversible thermal gelation in aqueous solutionsat a concentration of 20% or more (Schmolka, 1972). Thus, the polymersolution is a liquid at room temperature but gels rapidly in the body.Although the polymer is not degraded by the body, the gels dissolveslowly and the polymer is eventually cleared. This polymer has beenutilized in protein delivery (Morikawa et al., 1987; Jushasz et al.,1989) and skin burn treatments (Pautian et al., 1993).

e. PGA-PEO Hydrogels

Although PGA is not water soluble, bioerodible hydrogels based onphotopolymerized PGA-PEO copolymers have been synthesized and theirbiological activities investigated. (Sawhney et al., 1993; Sawhney etal., 1994; Hill-West et al., 1994). Macromonomers having a poly(ethyleneglycol) central block, extended with oligomers of α-hydroxy acids (e.g.,oligo(dl-lactic acid) or oligo(glycolic acid)) and terminated withacrylate groups were synthesized. These hydrogels were designed to formdirect contacts with tissues or proteins following photopolymerization,and act as a barrier.

These gels degrade upon hydrolysis of the oligo(α-hydroxy acid) regionsinto poly(ethylene glycol), the α-hydroxy acid, and oligo(acrylic acid).The degradation rate of these gels could be tailored from less than 1day to 4 months by appropriate choice of the oligo(α-hydroxy acid). Themacromonomer could be polymerized using non-toxic photoinitiators withvisible light without excess heating or local toxicity. The hydrogelspolymerized in contact with tissue adhere tightly to the underlyingtissue. In contrast, the gels were nonadhesive if they were polymerizedprior to contact with tissue. These hydrogels have been utilized inanimal models to prevent post-surgical adhesion and thrombosis of bloodvessels and initimal thickening following balloon catheterization.

Exogenous Factors Added to Tissue Scaffolds or Hydrogels

One aspect of the invention incorporates bioactive exogenous factorsinto the tissue scaffolds or hydrogels to stimulate tissue growth,relieve pain, fight infection, reduce inflammation or otherwisefacilitate the process of tooth repair in the methods of the invention.

a. ECM Components

In certain practices of the invention, growth factors that affect theproliferation of cells and tissues may be used in conjunction with thetissue scaffolds or hydrogels. It is preferable that the tissue-specificfunction of the proliferating cells that infiltrate the tissue scaffoldbe maintained. The function of the proliferating cells is stronglydependent on the presence of specific growth factors and ECM molecules(Stoker et al., 1990). For example, in vitro, it is known that cells canbe switched from a phase of tissue-specific gene expression to one ofproliferation simply by altering the ECM presentation to the cell(Mooney at al., 1992). Accordingly ECM proteins, hyaluronic acid orother components of the ECM may be incorporated into the tissuescaffolds used in the invention. If desired, any one of a variety oftissue scaffolds that incorporate specific ECM molecules may be used tosupplement the correct signalling to the host's proliferating cells(Green at al., 1979; Yannas et al., 1981; Bell et al., 1981; Stern etal., 1990; Compton et al., 1989; Dixit, 1994; Kasai et al., 1994;Cavallaro et al., 1994; Anselme, 1992; Koide et al., 1993; Guenard etal., 1992).

Synthetic materials that incorporate specific peptides to enhance celladhesion (McGrath et al., 1992; Barrera et al., 1993; Hubbell, 1993) maybe used, including those that incorporate a variety of differentpeptides in order to mimic the multi-functional nature of ECM molecules(Hynes, 1990). Growth factors promoting tissue development may belacking or deficient in the host tissue site that the engineered tissueis applied to.

b. Growth Factors

The use of growth factors in the context of cell proliferation andculture is generally well known in the art, although growth factors,other than those naturally present in the serum at low levels, have notbeen used in conjunction with oral tissue regeneration on a structuralmatrix in vivo. However, in that growth factors are routinely used inother contexts, one of skill in the art will readily understand how toapply growth factors in the context of the present invention based on ofinstant disclosure.

In general terms, it will be understood that a growth factor that hasalready been established to have a beneficial physiological effect on aparticular cell type should be chosen for use in regenerating tissuecontaining such cells. Certain growth factors may be used to stimulatethe proliferation of a wide number of cell types, whereas other growthfactors may have a more limited or defined cell-specificity.

Platelet-derived growth factor, (PDGF, e.g., PDGF-BB), which is onemember of the TGF supergene family of growth factors, may be used eitheralone or in combination with dexamethasone; or other growth factors.Particular examples of suitable growth factors include other members ofthe TGF supergene family, such as, BMP-2, BMP 4, BMP-7, VEGF, FGF-1,FGF-2, IGF-1, IGF-2, GDF-1, GDF-2, GDF-2, GDF-3, GDF-4, GDF-5, orcombinations of the same. PDGF-BB and dexamethasone are effective forthe growth of pulp, periodontal ligament and gingival fibroblasts(Rutherford et al., 1992a, 1992b; 1993a), and are particularly proposedfor use in connection with these aspects of the invention. U.S. Pat. No.5,149,691, incorporated herein by reference, describes the use ofcombinations of PDGF and dexamethasone for the repair and regenerationof tissues in vivo. U.S. Pat. Nos. 5,376,636 and 5,149,691, eachincorporated herein by reference, also describe the use of PDGF andglucocorticoids in tissue regeneration. Any such teachings may be usedin connection with the present invention. It is also known that thiscombination and PDGF/IGF-1 induce regeneration of the periodontium in ananimal model of periodontitis (Rutherford et al., 1992b; 1993a).

Bone Morphogenic proteins (BMP) such as those described in U.S. Pat.Nos. 4,795,804; 4,877,864; 4,968,590; 5,011,691; 5,013,649; 5,106,748;5,108,753; 5,116,738; 5,141,905; 5,166,058 and 5,187,076 are employed incertain aspects of the present invention. It has been demonstrated thata single application of BMP-7 to a freshly and partially amputateddental pulp induced reparative dentinogenesis in ferrets, monkeys andhumans (Rutherford et al., 1993b, 1994, 1995). Additionally, it has beendemonstrated that BMP-7 induced bone when implanted in gingiva,indicating that gingiva possess cells that are capable of formingmineralized tissue such as bone.

Direct local application of recombinant growth factors (e.g., BMP-2) hasbeen shown to induce reparative dentinogenesis in dogs and primates whenplaced on partially amputated dental pulps (Rutherford et. al., 1993b;Nakashima, 1994; Rutherford et. al., 1994), or on a freshly cut dentalsurface (“transdentinal” application; Rutherford et. al., 1995).However, in many clinical situations no pulp remains to stimulate.Moreover, direct application of the growth factors was not effectivewhen the pulp was inflamed. The present invention provides methods oftreatment where some pulp remains, whether or not it is inflamed, basedon recognizing that BMP-7 (alone or in combination with other growthfactors, such as BMP-2, BMP-4, BMP-7 or GDF-5) can be delivered from theimplanted tissue scaffold, or from a hydrogel in contact with the same,over time as the tissue scaffold or hydrogel is degraded in vivo. Thus,the problem of inflammation, which is commonly associated with mostdental conditions where the pulp is exposed, is not a barrier tostimulating growth of cells from the remaining pulp in the presentinvention, because the tissue scaffold is placed in contact with thepulp for a sufficient time for the inflammation to recede and/oroptionally, anti-inflammatory agents are combined with the tissuescaffold to facilitate the process.

The growth factors or stimulatory agents that are useful in the contextof the present invention may be purified from natural sources or may berecombinantly prepared proteins. They may be obtained from commercialsources, if desired. Those of skill in the art will know how to obtainand use such growth factors in the context of tissue regeneration inlight of the present disclosure.

c Anti-Inflammatory Agents

As previously mentioned, in certain practices an anti-inflammatory agentis combined with the tissue scaffold or hydrogel inserted into the toothof a subject. In one embodiment, the tissue scaffold or hydrogel mayinclude the anti-inflammatory agent alone, while in other embodiments itmay include the anti-inflammatory agent in combination with amorphogenic factor, antibiotic or other biologically active agent.Suitable anti-inflammatory agents include, amongst others, those in theclass of Cox-I and Cox II inhibitors. Examples of such agents includeacetyl-salicylic acid, acetaminophens, naproxen, ibuprofen and the like.Another example of a suitable class of anti-inflammatory agents includessoluble cytokine receptors such as Embrel™ or IL-1b binding receptors.The amount of anti-inflammatory agent used is adjusted so as to bereleased from the tissue scaffold or hydrogel over a period of about 2days or more.

d Analgesic Agents or Anesthetics.

In some embodiments, the tissue scaffold or hydrogels used in thepresent invention may include analgesic agents or anesthetics. Theanti-inflammatory agents mentioned above also serve as analgesic agents,thus analgesic agents include anti-inflammatories. In addition, theanalgesic agent may include local pain deadening agents (anesthetics)such as lidocaine, that provide local pain relief for a period of about30 minutes or more.

e Antibiotic Agents

In various embodiments, the tissue scaffold or hydrogels used in theinvention may include an antibiotic agent. There are numerous classes ofantibiotic agents suitable for the invention including, but not limitedto: tetracyclines, chemically modified tetracyclines, cyclosporins,those in the penicillin family, amoxicillan, gentamicin, erythromycin,chioramphenicol, florfenicol, vancomycin, everninomicin, cefotaxime,streptomycin, ciprofloxacin, nalidixic acid, bacitracin, enrofloxacin,and flavomycin.

In various embodiments, the tissue scaffold or hydrogels of theinvention may also include compositions for diffusing calcium phosphateions to assist in remineralization and repair of caries lesions.Suitable examples of such compositions are described, for example, inU.S. Pat. Nos. 5,833,954 and 5,993,786.

Amalgams, Cements and Fillers

The invention also includes sealing the hole in the tooth of the subjectwith a dental cement, filler or amalgam after placement of the dentalscaffold material and/or optional hydrogel. Suitable cements include forexample, any cement formulated for use as a base or liner, such as zincphosphate, glass ionomers or calcium phosphate. The cements may includea polymeribizable monomer, such as a carboxylated monomer. Oneparticular cement preferred for the invention also contains apolymeribizable monomer and di-calcium/tetra-calcium oxide as described,for example, in U.S. Pat. No. 6,398,859 and U.S. Pat. Pub. No.2002/0137812. The cement may also include flouride. One suitable exampleof a cement that contains flouride is described in U.S. Pat. No.6,056,930. Conventional metal based amalgams, although suitable in somepractices of the invention, should be avoided where possible.

Wafers

As mentioned in the definitions section, the tissue scaffolds used inthe present invention are shaped into wafers of predetermined size tocorrespond to the size of a hole drilled into the tooth of a subject.The wafers may be fabricated by any number of techniques and have avariety types of pore structures. FIG. 2 illustrates an exemplaryembodiment of a molded wafer 89, suitable for use in the invention. Inthe embodiment depicted in FIG. 2, the wafer 89 is formed into a spongewhere the interior is filled with the tissue scaffold material 82 ofsuitable porosity and pore size to permit the tissue to grow into andthrough out the matrix of the scaffold material 82.

The wafer 89 depicted in FIG. 2 is for example purposes only. The tissuescaffold material 82 wafers can be made into other forms suitable forthe purposes of the invention, including, but not limited to, openbarrel structures with an open lumenal space, concentric barrelstructures with concentric rings of the scaffold material 82 formingconcentric lumenal spaces, spiral barrel structures with a spirallumenal space, three dimensional interlinking sheets of the scaffoldmaterial 82, mats, spheres, cones and a variety of other geometric formsshaped to fit into a hole of corresponding size drilled into the toothof a subject.

Another aspect of the invention is wafers having particular markingsthereon to detect crushing. The highly porous tissue scaffolds used inthe invention are typically fragile structures that are subject toaccidental crushing when being manipulated or inserted into the hole ofcorresponding size drilled in the subject's tooth. As illustrated by theexamples in FIGS. 3A-3D, to readily detect crushing, the upper surface(or lower surface) of the wafers 89 of the invention may be marked witha pharmaceutically acceptable dye arranged in a pattern that that altersappearance when the wafer is crushed. In typical embodiments, thepattern includes regularly spaced markings 92 a-92 d so that crushing isdetected when an irregularity in appearance is exhibited. In theembodiment depicted in FIG. 3A, the markings 92 a are in concentriccircles. In the embodiment depicted in FIG. 3B the markings areregularly spaced hatch markings 92 b. In the embodiment depicted in FIG.3 c the markings are parallel lines 92 c. In the embodiment depicted inFIG. 3D the markings are regularly spaced dots or other geometric forms.The embodiments depicted in FIGS. 3A-3D are for example purposes only.Other markings or other means for detecting a crushed wafer can readilybe envisioned based on the present disclosure.

Tissue Scaffold Wafer Supports

Another aspect of the invention related to the prevention of crushing isa tissue scaffold support casing for supporting the tissue scaffoldwafers 89. Several embodiments of such support casings are depicted inFIGS. 4A-4D. The tissue scaffold wafers 89 made of the porous tissuescaffolding material 82 are supported by a casing comprised of supportelements 95, 97, 98, 99 101 and/or 103 made of a higher strengthpolymer. The higher strength polymer may be biodegradable or non-biodegradable. In one embodiment, the higher strength polymer is comprisedof PGA, PLA, PLLA, PLDLA, or PLGA having a porosity of less than 50% orless than 40% or less than 30% or less than 20%, or less than 10%. Thehigher strength polymer may also be non-porous (solid) in certainembodiments. Alternatively, the support material for the casing may bemade of a dental cement, amalgam, a fiber reinforced resin or anotherpolymeric material.

As illustrated in FIGS. 4A-4D, the casing support material is configuredto partially surround a portion the tissue scaffolding wafer 89.Generally, the tissue scaffolding wafer 89 is made of a material thathas a first crushing resistance and the casing support material isselected to have a second crushing resistance greater than the firstcrushing resistance. In the embodiment illustrated in FIG. 4A, thesupport casing material includes a horizontally disposed ring 97configured to contact the top and bottom of the tissue scaffold wafer89. In other embodiments, the ring 97 contacts one or the other of thetop and bottom surfaces of the tissue scaffold wafer 89. In all of theembodiments depicted in FIGS. 4A to 4D, the casing material includes atleast one columnar extension 95 extending from the an upper part of thecasing material downwardly along the sides of the wafer 89 between thetop bottom surfaces. However, other configurations of the support casingmaterial that contact the sides of the wafers, including for example,spirals, crosses, meshes and the like, may also be used in lieu of thecolumnar extensions.

In the embodiment depicted in FIG. 4B, instead of using the supportcasing ring 95, a circular pad 99 of the casing material is used on thetop surface. In this embodiment, the pad 99 provides an orientation forthe wafer, typically with the pad being on the top surface that willface the crown. The bottom surface will face the pulp 70 when positionedin a hole in subject's tooth. In the embodiment depicted in FIG. 4C, thecolumnar extensions are configured as brackets that include a laterallyextended portion 101 roughly perpendicular to the columnar extension 95so that at least one of the top or bottom of the wafer is positionedbeneath or above, respectively, the laterally extended portions 101. Inthe embodiment depicted in FIG. 4D, the laterally extended portion isformed as a brace 103 connecting at least two columnar extensions. Thebrace 103 extends across at least one of the top and bottom surfaces ofthe tissue scaffold wafer 89. Placement of the tissue scaffold wafer 89with the casing support elements 95, 97, 98, 99, 101, 103 in the toothof a subject is illustrated in FIG. 5. In a typical practice, the wafer89 with support casing elements 95, 97, 98, 99, 101, 103 is firstpositioned into the hole of corresponding size 105 drilled in the toothof a subject. The wafer 89 is secured into the hole with a base material106, which may for example be a hydrogel. The top of the tooth isfinally sealed with an appropriate amalgam, cement or other fillingmaterial 107.

Kits

FIG. 6 illustrates another aspect of the invention, which is a kit 110specifically designed to store a set of dry tissue scaffold wafers 89 ofa predetermined size to correspond with dental drilling bits, inconjunction with set of wetting solutions 112 to hydrate the tissuescaffold material 82. The kit includes a blister pack 114 carrying a setof individual dry wafers 89 arranged in an array of heights from about0.1 to about 0.8 mm, most typically about 0.5 mm in height, and indiameters of about 0.5 to about 5 mm, most typically about 2 to 5 mm indiameter. Each of the individual dry wafers 89 is positioned in a firstchamber 116 of the blister pack 114 that is positioned adjacent tosecond chamber 118 holding the wetting solution 112. The two chambersare separated by a thin wall 119 that is broken upon snapping a blistersnap 120 which causes the wetting solution 112 to flow into the firstchamber 116 and hydrate the wafer. The dry wafers 89 may includelyophilized or otherwise dried exogenous factors, morphogenic agents,analgesic agents, antibodies and the like. Alternatively, such exogenousfactors may be contained in hydrated form along with the wettingsolution 112.

In one embodiment, the kit 110 may also include a plurality ofpre-hydrated hydrogel plugs 106 of corresponding size arranged below thetissue scaffolds. The hydrogel plugs 106 may contain exogenous factors,morphogenic agents, analgesic agents, antibodies and the like alreadyloaded into the hydrogel 106. Alternatively, the kit may includecomponents for mixing the hydrogel 106 in batch for loading into asyringe or other suitable delivery device. The hydrogel components maybe mixed alone or in combination with the exogenous factors. In certainembodiments the exogenous factors may be included in a separate chamberfor mixing with the hydrogel components or the tissue scaffold wafers89.

The kits of the present invention thus typically include a means forcontaining the necessary components for use in close confinement forcommercial distribution. The kit may also include a variety of vials orother containers for holding the necessary components. Irrespective ofthe number or type of containers, the kits of the invention aretypically packaged with instructions for use of the kit components.

Vacuum Manipulator

FIG. 7 illustrates another aspect of the invention, which is a vacuummanipulator 130 configured for manipulating and placing the fragiletissue scaffolding materials in the hole of corresponding size drilledinto the tooth of a subject. The vacuum manipulator 130 includes anelongated vacuum tube 132 having a proximal end 134, a distal end 136,and walls 138 between the proximal and distal ends defining the vacuumtube 132. In certain embodiments, the vacuum tube 132 may have a bend140 along the length thereof to orient the proximal end 134 forinsertion into a mouth of subject along a plane defined by adjacenttooth crowns. The vacuum manipulator 130 further includes a suction cup142 attached to the proximal end 134. The suction cup 142 is configuredto be in fluid communication with the vacuum tube 132 and a vacuumsource 144. The suction cup 142 is sized to fit onto a surface of awafer 89 comprised of tissue scaffolding material 82. In certainembodiments particularly useful for wafers protected with a supportcasing, the outer perimeter of the suction cup 142 is configured to fitaround the ring 97 of support material so that the tissue scaffoldmaterial 82 is not contacted by the device.

The vacuum manipulator 130 may also include an in-line filter 146disposed in the vacuum tube 132 between the suction cup 142 and thedistal end 136. The in-line filter 146 has sufficient pore size to allowpassage of the gas fluid between the suction cup and the vacuum chamberwhile preventing the passage of bacteria or other pathogens. The vacuummanipulator 130 further includes a vacuum valve assembly 148 located atthe distal end 136 of the vacuum tube 132. The valve assembly 148 isoperable to close and open fluid access between the vacuum tube 132 andthe vacuum source 144 and, optionally, to open and close fluid accessbetween the vacuum tube 132 and a positive pressure source 150. Thepositive pressure source 150 provides a source of pressure that isgreater than the pressure of the vacuum drawn in the vacuum tube 132.The positive pressure source 150 may be coupled to the same vacuumsource 144, so that when the vacuum source is operated in an oppositedirection as used to draw the vacuum, a positive pressure flows into thevacuum tube. In alternative embodiments, the source of positive pressure150 may be an opening to ambient air pressure, or a second bulb, plungeror diaphragm. In yet another embodiment the source positive pressure 150may be an external compressor or in-house air source.

The valve assembly 148 is operated by manipulation of a control switch152, which is mechanically or electro-mechanically coupled to the valveassembly 148. The control switch 152 optionally has dial settings for aclosed position, a vacuum drawing position, and a vacuum release (orpositive pressure) position. In certain embodiments, the valve assembly148 or control switch 152 is adjustable to allow a controlled amount ofat least one of positive or negative pressure to be drawn or appliedwithin the vacuum tube 132.

In one embodiment, the vacuum source 144 may be a bulb, plunger or acompressible diaphragm. In another embodiment, the vacuum source 144 maybe an external aspirator or vacuum pump separate from the vacuummanipulator 130. In such embodiments, the vacuum manipulator 130 mayinclude a connector assembly for connecting the vacuum tube to theexternal vacuum source. In certain embodiments, the vacuum manipulator130 may optionally include a safety release configured with the valveassembly 148 to release any excessive vacuum drawn in the vacuum tube132.

In use, the operator positions the suction cup 142 over the top end ofthe wafer and manipulates the control switch 152 to draw a vacuum in thevacuum tube 132 thereby grabbing the top end of the wafer. The bottomend of the wafer is then positioned in place in the hole made in thesubject's tooth in contact with the pulp tissue 70. When properlypositioned, the operator releases the vacuum by operation of the controlswitch 152 thereby dislodging the wafer from the suction cup 142. Inembodiments configured with the positive pressure source 150, thecontrol switch is manipulated to a release position that appliespositive pressure into the vacuum tube 132 to assure the wafer isdislodged. The control switch 152, valve assembly 148, and positivepressure source 150 may be electromechanically configured tocooperatively operate to accomplish vacuum release and positive pressuresimultaneously or sequentially upon release of the vacuum.

Methods for Treating Dental Conditions

A. Treating Asymptomatic (“Nonsymptomatic”) Caries

Another aspect of the invention includes methods of treating dentalconditions using the wafers and devices of the present invention. Oneexample of such a condition treatable by the methods of the invention isasymptomatic caries. FIG. 8 illustrates the anatomy of a tooth afflictedwith asymptomatic caries that minimally invades the healthy coronalpulp. To treat asymptomatic caries, the decay/demineralized enamel anddentin is removed and the pulp exposure enlarged to a diameter of 0.5 to1.0 mm. A tissue scaffold material comprised of PLLA, PDLLA, PGA orPLGA, co-polymer, optionally patterned with calcium phosphate or calciumphosphate plus fluoride and shaped into a cylindrical wafer having adiameter to correspond with the diameter of the hole exposing the pulpas illustrated in FIG. 9. The wafer is hydrated in a physiologicallyacceptable fluid comprised, for example, of phosphate buffered saline,and is inserted into the hole so that the bottom end of the wafer is incontact with the pulp tissue in the corona) pulp.

A cement base or liner, for example, one comprised of a mixture of 60 to80% tetra-calcium phosphate and 20 to 40% di-calcium phosphate isapplied to cover the wafer and the upper recesses of the fill hole.Examples of suitable cement bases or liners include those described, forexample, in U.S. Pat. Nos. 6,398,859, 6,210,759, 6,206,959, 6,187,838,6,114,408, 6,001,897, or U.S. published application No. 2002/0137812.Alternatively, or in addition, a composition that promotesremineralization of the dentin that contains calcium phosphate with orwithout flouride, as described for example, in U.S. Pat. Nos. 5,037,639,5,268,167, 5,437,857, 5,460,803, 5,534,244, 5,562,895, 6,000,341, or6,056,930 may be added in the region of the dentin. The cement and/orremineralization composition is then covered with a standard permanentdental restorative material such as a composite resin or dental amalgam.After a sufficient period of time, typically about 1 week, cells fromthe exposed portion of the coronal pulp proliferate and infiltrate thetissue scaffold wafer and produce the components necessary to regeneratethe dentin. After a period of about 45-60 days, the tissue scaffoldwafer degrades leaving intact healthy dental tissue.

B. Replacing Root Canal Therapy

Still other conditions treatable by the methods of the invention arethose that typically require root canal therapy. Injury or infection ofadult dental pulp often necessitates such therapy. Root canal therapy iscommonly used in the case of severe caries where a substantial portionof the dentin and pulp tissue has been degraded, often into the vicinityof the root canal (FIG. 10).

Root canal therapy completely devitalizes a tooth and hence terminatesdentin formation and subsequent maturation. Unfortunately, the syntheticmaterials typically utilized to replace lost tooth structure are notcapable of completely replacing the function of the lost tissue, andoften fail over time. In a root canal procedure according to the priorart (FIG. 11), the entire pulp tissue and some of the dentin is removed.Gutta-percha is used to fill the void formed in the root canal andcoronal regions of the tooth, a resin or amalgam filling is then used toreplace the pulp, and often a metal or porcelain crown is cemented intoplace. Typically a root canal procedure leaves a devitalized tooth thatis prone to fracture and subsequent loss.

Rather than destroy all remaining pulp leaving a devitalized tooth, thepresent invention eliminates the need for root canal therapy whilepreserving the remaining pulp and regenerating new dentin.

FIG. 12 illustrates a procedure that replaces root canal therapyaccording to the present invention. A pulpotomy is performed whereby thecoronal pulp is removed to the level of the root canals. One or moretissue scaffold shaped into a cylindrical wafer having a correspondingdiameter of about 0.5 mm to 1 mm to correspond with the diameter of theorifice of the root canal is inserted to contact the pulp tissueremaining in the root canal.

In one practice, one or more of the wafers includes 1-100 pg of amorphogen such as a member of the TGF-β superfamily, for example, BMP-2,BMP-4, BMP-7 and/or GDF5. A cement base (FIG. 9), for example, onecomprised of a mixture of 60 to 80% tetra-calcium phosphate as mentionedabove, is applied to cover the wafer and the upper recesses of the fillhole. Alternatively, or in addition, a composition that promotesremineralization of the dentin that contains as mentioned above may beadded in the region of the dentin. The cement and/or remineralizingcomposition is then covered with a standard permanent dental restorativematerial such as a composite resin or dental amalgam. After a sufficientperiod of time, typically about one week, cells from the exposed pulptissue infiltrate the tissue scaffold wafers and produce the componentsthat regenerate the dentin.

Referring back to FIG. 13, in another practice of the foregoing methods,the morphogens are embedded in the hydrogel plug 108 that is appliedover the top of the tissue scaffold wafer prior to sealing the hole withthe cement and/or remineralization composition. In this practice, themorphogen gradually diffuses from the hydrogel into the tissue scaffoldmatrix over time.

In another practice, where very little or no healthy pulp tissue 70remains in the tooth, the tissue scaffold wafer inserted into any of theroot canal or coronal pulp chamber may be pre-seeded with dental pulpstem cells. Dental pulp stem cells may be obtained and identified, forexample, according to the methods described in PCT publication WO02/07679. These dental pulp stem cells are cultured in vitro and seededinto the tissue scaffold wafer according to the methods described, forexample, in U.S. Pat. Nos. 5,885,829, 6,281,256, in U.S. Pat.Publication No. 2002/0119180 A1 or by Young et al., (2002) J. Dent, Res.81 [10] p 695-700. A culture of about 5×10⁶ dental pulp stem cells willbe sufficient to regenerate the damaged pulp tissue and otherstructures.

C. Anti-Inflammatories and Antibiotics.

In certain practices of the forgoing methods, an anti-inflammatory agentand/or and antibiotic agent is included in the tissue scaffold wafer.The anti-inflammatory agent not only provides an analgesic effect in thevicinity of the tooth, but also promotes reduction of pulpitis. Asmentioned in the Background herein before, the morphogenic agent BMP-7,when directly applied as a moistened powder to the tooth was noteffective in promoting pulp growth in ferrets having pulpitis, but waseffective in promoting pulp growth in non-inflamed tissue. BMP-7 waseffective in stimulating pulp growth in ferrets having pulpitis whendelivered from cells transduced with a gene encoding BMP-7.

While not being bound by theory, it is believed that the differencebetween the effectiveness of directly applied BMP-7 versus ex-vivocellularly produced BMP-7 is due to the presence of inflammation and thetiming of delivery. Directly delivered BMP-7 remains active for only ashort period of time during which inflammation of the pulp persists. Itis believed that inflamed pulp tissue is not responsive to the activityof morphogens such as BMP-7. However, cellularly delivered BMP-7 iscontinuously produced over the time period required for the inflammationto recede, after which time the non-inflamed tissue become responsive tothe morphogen. The present invention eliminates the need to use ex-vivotransgenic cells to deliver the morphogen over time because themorphogen is embedded in the tissue scaffold wafer (or in a hydrogel)which gradually releases the morphogen over time as the wafer degrades.In addition, the presence of the anti-inflammatory agent in the waferpromotes more rapid recovery of the inflamed tissue, making itresponsive to the morphogen. Further, when an antibiotic is included,its presence not only reduces the risk of infection, but also reducesthe likelihood of an inflammatory response associated therewith.

As an alternative to embedding the anti-inflammatory agent and/or theanti-biotic in the tissue scaffold or hydrogel, the subject's tooth maybe treated by topical application of the agent or anti-biotic at one ormore sessions after the wafer has been implanted. This direct treatmenthas the benefit of permitting the therapist to precisely control theamount of anti-inflammatory agent or anti-biotic being applied to thetooth over time.

To aid the practitioner in making and using the tissue scaffold wafers89 that may be used in the practice of this invention, the followingExamples are provided as a guide to exemplify the manufacture, testingand modifications of the wafers used in the invention.

EXAMPLE I Fabricating Hollow Tube Tissue Scaffold Wafers of PLA, andPDLLA and PLGA

A. Materials

PLA, PDLLA, and the 85/15 and 50/50 PLGA were purchased from Medisorb(Cincinnati, Ohio), chloroform from Mallinckrodt (Paris, Ky.),polystyrene standards from Polysciences (Warrington, Pa.), aluminumbacked tape from Cole-Parmer (Chicago, Ill.), phosphate buffered salineand DMEM medium from Gibco (Grand Island, N.Y.), T_(max) film fromKodak, Lewis rats, 250 to 300 g, from Charles River (Wilmington, Mass.),and methoxyflurane from Pitman-Moore Inc. (Mundelein, III.).

Molecular weights of the various polymers were determined by gelpermeation chromatography (Perkin-Elmer, Series 10, Newton Centre,Mass.), using polystyrene standards to generate a calibration curve. PLAhad a molecular weight (M_(w)) of 74,000 (M_(w)/M_(n)=1.6); poly-(D,Llactic) acid had M_(w)=77,000 (M_(w)/M_(n)=1.8); 85/15 copolymer hadM_(w)=69,000 (M_(w)/M_(n)=1.9); 50/50 copolymer M_(w)=43,400(M_(w)/M_(n)=1.43). Differential scanning calorimetry was utilized toconfirm the amorphous nature of all of the polymers except PLLA, whichexhibited the expected crystallinity.

B. Device Fabrication

Hollow tubes were formed by a two-step process; porous films of thepolymers were first fabricated, and these films were then formed intohollow tubes. To fabricate porous films, the polymer was dissolved inchloroform to form a 1.56% solution (w/v). Eight ml of this solution wascast into a 5 cm glass petri dish covered with a sheet of aluminumbacked tape. Sieved sodium chloride crystals between 150 and 250 μm indiameter were dispersed evenly over the solution (0.375 g NaCl/dish),and the chloroform was allowed to evaporate at room temperature. Apolymer film with entrapped NaCl particles resulted.

The salt particles were leached out of the film by immersion in 800 mlof deionized water for 48 hr at 370° C. with constant shaking. The waterwas changed every 8 hr during the leaching period. This procedureyielded a highly porous, thin membrane. In one practice sections werecut from the resulting films (1.3.times.1.5 cm), with a razor blade, androlled around Teflon 15 cylinders with an outer diameter of 0.32 cm. Thesurfaces of the films that were adjacent to the aluminum backed tapewere always placed adjacent to the Teflon cylinder.

In another practice, Teflon cylinders having diameters that correspondto the diameter of holes to be drilled in the tooth of a subject areselected. The height of the cylinder is similarly selected to correspondto the depth of the portion of the hole drilled into the dentin layer.Typical selected diameters sizes are between about 2 to 5 mm. Typicalheights are about 1 to 3 mm. Sections of the membrane are cut from theresulting films to correspond to the diameter of the Teflon cylindersand leaving about 10% to 30% of overlapping material over the ends.

The overlapping ends of the film were sealed together by brieflyexposing one edge to chloroform, and manually pressing the overlappingends together. The chloroform temporarily dissolved the polymer on thesurface of each of the overlapping ends, and after the chloroformevaporated the overlapping ends were sealed together. The tubes werethen slipped off of the teflon template. The ends were closed by placinga circular piece of the same porous films over the ends, and sealing asabove with chloroform. Tubes 1.5 cm long, with an inner diameter of 0.32cm resulted. Tubes were lyophilized to remove residual solvent, andsterilized by exposure to ethylene oxide for 24 hr at room temperature.Tubular tissue scaffolds formed by this method are illustrated in FigureA.

In an alternative practice, tubular tissue scaffolds are formed with aconcentric tubular lumen. To form tissue scaffolds with concentriclumens, a set of tubular scaffolds varying in diameter by a selectedspacing distance are made according to the steps outlined above, except,that prior to sealing the ends, the set of tubular scaffolds areinserted into one another to share a common origin about a circle. Thisresults in an interior lumen having concentric luminal spaces asillustrated in Figure A1. After arranging the concentric tubes, the endsare sealed with a circular strip of the porous membrane as describedabove.

In another alternative practice tubular tissue scaffolds are formed witha spiral tubular lumen. Instead of wrapping the porous membrane around aTeflon cylinder, the membrane is positioned on top of Teflon tape havinga thickness of about 0.1 to about 0.3 mm. The Teflon tape and membranetogether are then rolled into a tubular shape of the desired diameter.Rolling in this fashion results in a spiral arrangement of the membranestrip with a separation distance across adjacent walls of the spiralequal to the thickness of the Teflon strip as depicted in Figure A2. Thebottom of the rolled strip is sealed with a circular membrane strip asbefore, the Teflon tape is removed and the top is sealed with anothercircular membrane, resulting in a closed tubular tissue scaffold with aspiral shaped lumen.

The pore structure of films formed with the particulate leachingtechnique can be controlled by varying the ratio of polymer to salt inthe film fabrication process. Films fabricated with a polymer/salt ratioof 1/3 exhibited large pores on the air surface, approximately the sizeof the salt particles utilized to form the pores. These films had muchsmaller pores on the surface of the film exposed to the teflontape-coated glass surface. These pores corresponded closely in structureto the salt particles utilized to create the device porosity.

Decreasing the ratio of polymer to salt from 1/3 to 1/24 resulted in theformation of larger pores on the air surface of the films, and thesepores were not as uniform as the pores formed at lower salt loadings.Larger pores also formed on the film surface exposed to the glass dishas the polymer to salt ratio was decreased. At a very low ratio (1/12)both sides of the films exhibited a similar pore structure.

Films fabricated from all polymers (PLLA, PDLLA, 50/50 PLGA, and 85/15PLGA) could be readily formed into hollow tubes. There was nosignificant differences in the pliability of films formed from thedifferent polymers. However, films formed using a polymer/salt ratiolower than 1/12 were quite brittle, and the manipulation needed to formtubes from these films often resulted in their fracture.

Films fabricated using a polymer/salt ratio of 113 were utilized tofabricate the tubular devices used in all subsequent studies. Thesefilms had a thickness of 320±50 μm (mean.±. sd, n=36), a porosity of87±0.4% (n=12), and a volume average pore diameter of 150±50 μm (n=12).

Devices fabricated from all four polymers resisted compression bymoderate forces (50 mN), but the ability of devices to withstand alarger compressional force (150 mN) was dependent on the polymer used infabrication. PLLA tubes exhibiting the least compression under thisforce, and 50/50 PLGA devices exhibiting the greatest. The compressionwas viscoelastic in all cases, as the devices only partiallydecompressed after the force was removed. While devices fabricated fromPLLA and PDLLA consistently resisted this load, PLGA devices did notexhibit a consistent response. The 50/50 PLGA tubes showed the greatestvariability, as testing of multiple tubes gave widely varying results.Devices which compressed greater than 50% after force applicationtypically showed little elastic recoil after the force was removed.Thus, devices which compressed more than this were considered to fail atthis loading, and the failure rates of devices fabricated from thevarious polymers is given in Table 7. TABLE 7 In vitro and In vivoCompression Test Failure of Devices In vitro In vivo Number of Number ofPolymer Sample Polymer Samples Polymer Samples Failure Rates SamplesFailure Rates PLLA 6  0% 2  0% PDLLA 5 20% 6  0% 85/15 5 40% 9 44% PLGA50/50 16 58% 6 50% PLGAA. Compression testing was performed with a force of 150 mN. A devicewas considered to fail if it compressed to less than 50% of the originaldiameter.B. Devices were implanted into the mesentery or omentum of Lewis ratsfor 7-28 days. Histological sections were examined, and devices wereconsidered to fail if they compressed and did not maintain theiroriginal, tubular shape.

Different erosion times may be required of devices utilized to engineervarious tissues. The erosion kinetics of devices fabricated in thisexample was governed by the polymer utilized to fabricate the devices.The time for complete erosion could be varied between 10 weeks and overa year by varying which polymer was utilized to fabricate the device.

The compression resistance of the tubular devices fabricated in thismanner were dependent on the polymer utilized to fabricate the devices.Devices fabricated from both PLLA and PDLLA resisted compressionalforces and maintained their structure both in vitro and in vivo. Devicesfabricated from PLGA did not resist compressional forces as well eitherin vitro or in vivo. The compression resistance of tubes fabricated from50/50 PLGA was inconsistent at high compression forces (150 mN).

The tubular devices and tissues are only exemplary of how a polymermatrix can be used to provide a 3-D structure required to match a nativetissue structure. Particular devices optimized for oral tissues can thusbe generated.

EXAMPLE II Making PGA Tubular Tissue Scaffolds Stabilized by SprayCasting with PLLA and PLGA

Another method to stabilize PGA meshes, described in this example, is toatomize solutions of poly(L-lactic acid) (PLLA) and a 50/50 copolymer ofpoly(D,L-lactic-co-glycolic acid) (PLGA) dissolved in chloroform and tospray over meshes formed into hollow tubes. The PLLA and PLGA coated thePGA fibers and physically bonded adjacent fibers. The pattern and extentof bonding was controlled by the concentration of polymer in theatomized solution, and the total mass of polymer sprayed on the device.The compression resistance of devices increased with the extent ofbonding, and PLLA bonded tubes resisted larger compressive forces thanPLGA bonded tubes. Tubes bonded with PLLA degraded more slowly thandevices bonded with PLGA.

PGA fiber meshes are stabilized by physically bonding adjacent fibresusing a spray casting method. Poly L-lactic acid (PLLA) or a 50/50copolymer of lactic and glycolic acid (PLGA) was dissolved inchloroform, atomized, and sprayed over a PGA mesh formed into a tubularstructure. Following solvent evaporation, a physically bonded structureresulted, and the pattern and extent of PGA fiber bonding was controlledby the processing conditions.

These tubular devices were capable of withstanding large compressiveforces in vitro (50-200 mN), and maintained their structure in vivo. Thespecific mechanical stability was dictated by the extent of physicalbonding and the polymer utilized to bond the PGA fibers.

A. Tube Fabrication and Characterization

Rectangles (1.3.times.3.0 cm) of the non-woven mesh of PGA fibers werewrapped around a teflon cylinder (o.d.=3.0 mm) to form a tube, and thetwo overlapping ends were manually interlocked to form a seam. Theteflon cylinders were then rotated at 20 rpm using a stirrer (Caframo;Wiarton, Ontario, Canada). Solutions of PLLA and PLGA dissolved inchloroform (1-15% w:v) were placed in a dental atomizer (DevilbusCorp.), and sprayed over the rotating PGA mesh from a distance of 6inches using a nitrogen stream (18 psi) to atomize the polymer solution.

The PLGA and PLLA had molecular weights (Mw) of 43,400 (Mw/Mn=1.43) and74,100 (Mw/Mn=1.64), respectively. Molecular weights were determined bygel permeation chromatography, as described above.

While PLLA and copolymers of lactic and glycolic acid are soluble inchloroform, PGA is very weakly soluble in this solvent. Thus, the PGAfibers are largely unchanged by this process. After spraying wascompleted, the tubes were lyophilized to remove residual solvent,removed from the teflon cylinder, and cut into specific lengths. Thetubes were sterilized by exposure to ethylene oxide for 24 hr, followedby degassing for 24 hr.

The mass of PLLA and PLGA that bonded to the PGA scaffolds wasdetermined by weighing PGA devices before and after spraying. Electronmicroscopy, mechanical analysis and erosion characteristics were asdescribed above.

1. Bonding Tubes with PLLA

To determine whether PGA scaffolds could be stabilized by physicallybonding adjacent fibers, chloroform containing dissolved PLLA (1-15%w:v) was sprayed over the exterior surface after the PGA mesh waswrapped around a teflon cylinder to form a tube. The PLLA formed acoating over the exterior PGA fibres after the solvent evaporated, andphysically bonded adjacent fibres. The tubes formed in this manner couldbe easily removed from the teflon cylinder for characterization and use.

The pattern of bonding was controlled by the concentration of the PLLAin the atomized solution, even though the time of spraying was adjustedto maintain an approximately constant mass of PLLA on the devices underthe various conditions. Spraying with a solution containing 1 or 5% PLLAresulted in extensive bonding of PGA fibres without significantlyblocking the pores of the PGA mesh. Spraying with a 10% solution of PLLAalso bonded fibers, but resulted in the formation of a PLLA film on theexterior surface of the PGA mesh that contained only small pores.Spraying with a solution containing 15% PLLA had a similar effect,although the polymer film that formed was less organized. In all cases,the PLLA coated and bonded fibers only on the exterior surface of thePGA mesh, as no coating or bonding of fibers was observed on theinterior surface of the PGA mesh.

The compression resistance of bonded tubes was assessed in vitro todetermine which patterns of bonding resulted in the most stable devices.Unbonded tubes were completely crushed by a force of 5 mN, but bondedtubes were capable of resisting forces in excess of 200 mN. However, theability of bonded tubes to resist a given compressional force wasdependent on the pattern of bonding. For example, tubes bonded with 1 or15% PLLA were significantly compressed by a force of 200 mN, while tubesbonded with a solution of 5 or 10% PLLA were only slightly compressed bythis force. The compression was viscoelastic in all cases, as thedevices only partially decompressed after the force was removed. Uniformproperties were observed with respect to the position along and around atube.

To determine if the extent, as well as the pattern, of bonding couldvary the compression resistance of tubes, an atomized dispersion of 5%PLLA was then sprayed over the devices for different times. Lengtheningthe spraying time from 10 to 60 seconds increased the mass of PLLA onthe devices. Infrequent bonds between adjacent fibers resulted fromspraying for 10 seconds. Spraying for more extended periods increasedthe PLLA coating over the PGA fibers, and the extent of bonding.

The ability of these tubes to resist compressional forces and maintaintheir shape was quantitated again using thermal mechanical analysis. Thecompression resistance strongly depended on the extent of bonding, astubes that were more extensively bonded had a greater resistance todeformation. The compression that did occur under these conditions wasagain a combination of a reversible, elastic strain, and an irreversibledeformation. Some tubes were also exposed to an aqueous environmentbefore testing to determine whether this environment for 24 hr woulddestabilize the tubes. The aqueous environment had a slight, detrimentaleffect on the stability of bonded tubes, but they were still capable ofresisting large compressive forces.

2. Bonding Tubes with PLGA

To determine whether this technique of stabilizing PGA devices could beutilized with a variety of polymers, the previous study was repeatedusing a 50/50 copolymer of lactic and glycolic acid. The mass of polymerbonded to the devices and the extent of physical bonding was againregulated by the time an atomized dispersion of the bonding polymer wassprayed over the PGA fibers. Once again, bonding increased thecompression resistance of devices formed into a tubular structure.However, these devices were not able to resist the same compressionalforces as PLLA bonded devices.

Tubes bonded with PLLA were capable of resisting forces up to 200 mN,while tubes bonded with PLGA were only capable of resisting forcesslightly greater than 50 mN: The difference between devices stabilizedwith PLLA and PLGA was even more striking when the devices were testedafter immersion in phosphate buffered saline for 24 hr. PLGA bondedtubes, in contrast to PLLA bonded tubes, were significantly weakened bythis treatment.

The bonding approaches described herein permit a variety of bondingpolymers to be utilized, and allows the fabrication of variousthree-dimensional scaffolds. It also results in bonding only of theoutermost fibres of the device in contrast to other methods. Thispreserves the desirable features of the PGA mesh (high porosity, highsurface area/polymer mass ratio) throughout the interior sections.

This approach also allows both the extent and pattern of bonding to beeasily controlled. Extensive coating and bonding of fibers resulted whenthe polymer concentration in the atomized solution was low (1-5%).Increasing the concentration of polymer in the atomized solution to 10%resulted in the formation of a relatively smooth film over the externalsurface of PGA meshes, and utilizing a 15% solution resulted in theformation of a fibrous, nonhomogeneous film over the PGA meshes.Increasing the polymer concentration raises the viscosity of thissolution, and this likely increases the droplet size which is formedduring the atomization process. This will effect how these dropletspenetrate the PGA mesh, how they aggregate on the PGA mesh, and the rateof solvent evaporation. All of these factors will effect the pattern ofbonding.

For delivery into a hole in a subject's tooth, the tissue deliverydevice should preferably maintain a pre-configured geometry in the faceof external forces during the process of tissue development. Themagnitude of the compressive forces that are exerted on implanteddevices by the surrounding tissue are unclear and will vary depending onthe implant site. The magnitude of forces utilized in the presentexample to quantitate the compression resistance of devices in vitro was50-200 mN. This results in pressures ranging from approximately 50-200mm Hg (assuming complete and continuous contact between the TMAcompression tip and the tube). These pressures are in the same rangeobserved in blood vessels.

EXAMPLE III Making a PLGA Sponge Matrix Tissue Scaffold Wafer

Pellets of an 85:15 copolymer of D,L-lactide and glycolide (PLGA) waspurchased from Boehringer Ingelheim (Henley, Montvale, N.J., USA), andutilized to fabricate polymer matrices in all experiments. The intrinsicviscosity of the polymer was about 1.3-1.7. Polymer pellets were groundusing a Tekmar grinder (Bel-Art Products, Pequannock, N.J., USA), andsieved to obtain particles ranging from 106 to 250 pm. In certainexperiments the polymer particles were mixed with sodium chlorideparticles (Mallinckrodt, Paris, Ky., USA). The salt particles weresieved to yield a range of sizes, and the weight ratio of NaCl:PLGAmasses ranged from 0 to 50. In all cases, the total mass of PLGA andNaCl was held constant at 0.8 g. The mixtures of PLGA and NaCl wereloaded into a KBr die (1.35 cm in diameter; Aldrich Chemical Co.,Milwaukee, Wis., USA), and compressed at 1500 psi for 1 minute using aCarver Laboratory Press (Fred S. Carver, Inc., Menominee Falls, Wis.,USA) to yield solid discs (thickness=3.4 mm). The samples were thenexposed to high pressure CO₂ gas (800 psi) for 48 hours to saturate thepolymer with gas. A thermodynamic instability was then created bydecreasing the gas pressure to ambient pressure. This led to thenucleation and growth of CO₂ pores within the polymer matrices. The NaClparticles were subsequently removed from the matrices by leaching thematrices in ddH₂O for 48 hours. All processing steps were performed atambient temperature.

Porous sponges were also fabricated using a previously described solventcasting-particulate leaching technique. (A. G. Mikos, A. J. Thorsen, L.A. Czerwonka, Y. Bao, and R. Langer, “Preparation and characterizationof poly(L-lactic acid) foams,” Polymer, 35, 1068-1077 (1994).) In thisprocess, PLGA was dissolved in chloroform (Mallinckrodt; Paris, Ky.,USA) to yield a solution of 10% (w:v), and 0.12 ml of this solution wasloaded into Teflon cylinders (diameter 0.5 cm; Cole Parmer) packed with0.4 g of sodium chloride particles sieved to a size between 250 and 500mm. Following solvent evaporation, polymer films with entrapped saltparticles (3 mm thick) were carefully removed from the molds. The saltwas removed by immersing films in distilled water for 48 hrs.

The porosity of samples was initially determined by gross measurementsand weights after processing using the following equation:porosity(%)=1−[(weight/volume)/(density of polymer)].times.100  Eqn. 1

Thermogravimetric analysis was utilized to determine the amount of saltresidue that remained in the sponge after leaching. Matrices were heatedfrom 150° C. to 300° C. at a constant rate of 10° C./min, and theresidual mass was monitored.

Photomicrographs showed that gas foaming, alone, of solid polymer discsled to the formation of highly porous matrices. However, these matriceshad a nonporous skin on the external surfaces and the pores were largelyclosed, as expected from previous studies. (D. J. Mooney, D. F. Baldwin,N. P. Suh, J. P. Vacanti, and R. Langer, “Novel approach to fabricateporous sponges of poly(D,L-lactic-coglycolic acid) without the use oforganic solvents,” Biomaterials, 17, 1417-1422 (1996).) In contrast,gas-foaming and subsequent leaching of discs containing a highpercentage (95%) of large (250<d<425 μm) NaCl particles, according tothe invention, led to the formation of highly porous, open pore matriceswith no evidence of an external, non-porous skin. The pore structureobserved in cross-sections of these matrices was similar to thatobserved in cross-sections of matrices formed with a SC/PL technique.However, the pore structure of matrices formed from the SC/PL process isoften not uniform throughout the matrix due to evaporation of theorganic solvent and subsequent increase in the polymer concentration ofthe remaining solution entrapped within the salt bed. For example, thesurface of these matrices that is adjacent to the glass coverslip duringprocessing is shown in photomicrographs to be typically less porous thanthe remainder of the matrix. In contrast, the pore structure of gasfoamed-particulate leached (GF/PL) matrices was uniform throughout thematrix and on the exterior surfaces. TGA analysis of matrices indicatedthat negligible amounts of NaCl remained after leaching. There was atrace of a white residue left in the dish. To confirm that the gasfoaming was responsible for the formation of stable matrices, controlsamples were compression molded, but not foamed. Leaching of the NaClfrom these matrices led to complete breakdown of the matrices.

The ratio of NaCl:PLGA and the size of NaCl particles in GF/PL matriceswere varied to determine the range of porosity and pore structure thatcould be obtained with this process. The gross porosity of thesematrices increased from 85.1%.±.2.3 to 96.5%.±.0.5 as the ratio ofNaCl:PLGA was similarly increased. At constant NaCl (95%), the increasein salt particle diameter had very little effect on the overallporosity. However, photomicrographs showed that as the salt diameter wasincreased, the pore size increased in parallel.

The stability of the matrices was assessed using compressive and tensilemechanical tests. In general, the GF/PL matrices exhibited improvedmechanical properties as compared to the SC/PL matrices. The averagecompression moduli were 159.±.130 kpa and 289.±.25 kPa for the SC/PL andGF/PL matrices, respectively. The average tensile moduli were 334.±.52kPa for the SCIPL matrices and 1100.±.236 kPa for the GF/PL matrices(Table 8). This data represents a 80% increase in compression strengthand a 300% increase in tensile strength. TABLE 8 Gross porosity ofsponges. NaCl Concentration (%) Diameter (μm) 106-250 250-425 >425 80 —85.1 ± 2.3 — 90 87.3 ± 1.9 91.5 ± 1.4 — 95 93.9 ± 0.9 94.6 ± 0.9 95.0 ±0.8 97 — 96.5 ± 0.5 —

EXAMPLE IV Making Tissue Scaffold Matrix Sponges by Foaming CastedPolymer Disks

Pellets of poly L-lactic acid [PLLA], a 50:50 copolymer of D,L-lactideand glycolide (50:50 PLGA) with intrinsic viscosity (i.v. of 0.2 dL/g),a 75:25 PLGA copolymer (i.v.=1.3), and an 85:15 PLGA copolymer(i.v.=1.4) were obtained from Boehringer Ingelheim (Henley, Montvale,N.J., USA). PGA, 50:50 PLGA (i.v.=0.8) and 85:15 PLGA (iv=0.63) werepurchased from Medisorb (Cincinnati, Ohio, USA). 85:15 PLGA (i.v.=3.63)was obtained from Purasorb (Lincolnshire, Ill., USA).

The solid polymer (PLLA, PLGA, PGA) was ground (after freezing withliquid nitrogen) using a Scienceware Micro-Mill (Bel-Art Products,Pequannock, N.J., USA) and sieved to a diameter of 106-250 5 μm. NaCl,obtained from Fisher Scientific (Pittsburgh, Pa., USA), was sieved to adiameter of 250-425 5 μm for use in certain experiments. Solid polymerdisks were formed by placing 150 mg polymer (PGA, 50:50 PLGA, 75:25PLGA, 85:15 PLGA, and PLLA) into a round stainless steel KBr die withdiameter 1.35 cm (Aldrich Chemical Co., Milwaukee, Wis., USA) andcompressing for 60 seconds at 1500 psi in a Carver Laboratory Press(Fred S. Carver, Inc., Menominee Falls, Wis., USA). This method yieldssolid disks to be foamed. All samples were fabricated in triplicate.

The disks were foamed in a high pressure vessel using C0₂, N₂, or He at850 psi. After the disks were equilibrated (148 hours) with the gas, thepressure was reduced to ambient. The resulting thermodynamic instabilitycaused nucleation and growth of gas pores within the polymer matrix.85:15 solid polymer disks (i.v.=1.4) were foamed for 1 hour in CO₂ andthe pressure was released at different rates (1, 2.5, 5, 10 minutes) todetermine if the rate of pressure release affects the final structure ofthe sponges. All processing steps were performed at ambient temperature.

Polymer/NaCl disks were fabricated in a similar way using 40 mg polymerand 760 mg NaCl, compressed into disks. Following foaming, the diskswere placed in distilled water in order to remove the NaCl. Thisleaching solution was changed several times over the course of about 18hours. The disks were considered to be completely leeched when theleeching solution did not give a precipitate with AgNO3. If Cl— ispresent in solution, it precipitates with Ag+ to form a whiteprecipitate. The failure of this precipitate to form indicated that theNaCl is completely removed from the scaffolds. The disks were then airdried overnight, measured and weighed, and stored in a desiccator undervacuum. The polymer disks were measured and weighed immediatelyfollowing foaming, then stored in a desiccator under vacuum.

In order to calculate the porosity of the foamed disks, a boley gaugewas used to measure the diameter and thickness of each disk. The diskswere weighed on a Mettler balance and the following equation was used:(d=polymer density, g=disk wt, cm3=calculated disk volume).porosity=100[1−(g/cm3)/d]

a Foaming Solid Polymer Disks

In the first series of experiments, solid polymer disks were foamed toinvestigate the role of the gas type, pressure release rate, and polymercomposition and molecular weight on the porosity of polymer matrices.85:15 PLGA matrices were foamed for 1 hour with several different gases(CO₂, N₂, He). Significant porosity resulted from foaming with CO₂ ascompared to N₂ and He. The “prefoam” porosity refers to the calculatedporosity following disk preparation, but prior to high pressureequilibration. Visualization of matrices foamed with CO₂ revealed ahighly porous matrix consisting largely of closed pores.

In the next study, the rate of release of pressure was varied from 1 to10 minutes total time. The porosity of the matrices was relativelyconstant regardless of pressure release rate, except in the case of avery rapid release, when the gas froze within the chamber. This led to asmall decrease in the matrix porosity.

The effect of the polymer composition was investigated by usingdifferent copolymer ratios of PLGA (pure PGA, 50:50, 75:25, 85:15 PLGAand pure PLLA). Neither PGA nor PLLA foamed appreciably. The copolymersall foamed to a porosity greater than 90%. In fact, the 75:25 copolymerfoamed so extensively that it did not maintain its integrity in thepressure release/gas expansion phase and literally fell apart. Hence, noporosity value could be calculated for that sample.

In order to study the effect of polymer molecular weight on poreformation, disks of 85:15 PLGA with intrinsic viscosity (i.v.) rangingfrom 0.63 to 3.59 dL/g were foamed in 850 psi CO₂ for 24 hours with apressure release of 2.5 minutes. The high i.v. PLGA led to matrices withrelatively low porosity, whereas the lower i.v. PLGA resulted in muchhigher porosity.

b Foaming Polymer/NaCl Disks

In the second series of experiments, NaCl was incorporated into thepolymer disk for the purpose of creating an open pore structure.Different variables (equilibration time and polymer composition) werestudied in order to determine their effects on the structure andstability of the scaffolds. The results of the first series ofexperiments led us to use CO₂ as the foaming gas, and a pressure releasetime of 2.5 minutes in this series of experiments. Examination of atypical matrix formed by foaming 85:15 PLGA with NaCl in CO₂ shows ahighly porous structure with largely open, interconnected pores.

In the first study, the equilibration time was varied from 1 to 48hours. The porosity of the matrices was relatively constant forequilibration times greater than 6 hours, but decreased forequilibration times under 6 hours. Matrices fabricated with variousequilibration times were subsequently tested to determine if theequilibration time affected their mechanical properties. Even thoughmaximal porosity was achieved with 6 hours of gas equilibration, astronger scaffold was produced with longer equilibration times.

The polymer Composition was next varied to determine if results similarto those in the first series of experiments would be obtained.Copolymers of PLGA led to a much greater porosity than did thehomopolymers PGA and PLLA. Both the PLLA and PGA disks disintegrated inthe leaching process, indicating that little, if any, foaming hadoccurred. Even though all PLGA copolymers led to matrices with similarporosities, the matrices fabricated from PLGA with higher lactic acidcontent were more rigid.

EXAMPLE V Tissue Scaffold Sponge Matrix Wafers Made of PLA or PLGA—withPVA

The present example concerns sponges fabricated from poly-L-lactic acid(PLA) infiltrated with polyvinyl alcohol (PVA). Highly porous sponges(porosity-90-95%) were fabricated from PLA using a particulate leachingtechnique. To enable even and efficient cell infiltration, the deviceswere infiltrated with the hydrophilic polymer polyvinyl alcohol (PVA).

A. Sponge Fabrication and Characterization

Porous sponges were formed from PLA and an 85/15 copolymer of D,L lacticacid and glycolic acid (85/15 PLGA) (Medisorb; Cincinnati, Ohio) using avariation of a previously described particulate leaching technique(Mikos et al., 1994). The PLA and 85/15 PLGA had molecular weights(M_(w)) of 74,000 (M_(w)/M_(n),=1.6), and 69,000 (M_(w)/M_(n)=1.9),respectively. Molecular weight determination was performed using gelpermeation chromatography with polystyrene molecular weight standards.

The polymers were dissolved in chloroform (Mallinkrodt; Paris, Ky.) toyield a solution ranging from 10-20% (w:v), and 0.12 ml of this solutionwas loaded into Teflon cylinders (diameter=21.5 cm, height=25 mm; ColeParmer) packed with 0.4 g of sodium chloride particles sieved to a sizebetween 250 and 500 μm. Following solvent evaporation, polymer filmswith entrapped salt particles (1 mm thick) were carefully removed fromthe molds. The salt was removed by immersing films in distilled waterfor 48 hr. The water bath was changed 3 times daily. The volume ofpolymer solution and salt mass loading were linearly increased tofabricate thicker sponges.

To infiltrate sponges with PVA (Aldrich Chem. Co.; Milwaukee, Wis.; MW3000, 75% hydrolyzed) or the Pluronic F 108 surfactant (BASF;Parsippany, N.J.), sponges were immersed for 16 hr in an aqueoussolution containing 1-100 mg/mL of PVA or Pluronic in phosphate bufferedsaline (PBS). The sponges were subsequently removed from the solution,dried, and lyophilized. The mass of devices before and after coating wasquantitated to determine the mass of incorporated PVA or surfactant. Todetermine whether the incorporated PVA was permanently associated withthe sponges, some sponges were subsequently soaked in a solution of PBSovernight, air dried at room temperature, lyophilized, and reweighed.All sponges were sterilized before use by exposure to ethylene oxide.

The ability of aqueous solutions to wet the sponges was determined byplacing a volume of distilled water equivalent to the void volume of thesponge (as determined using mercury porosimetry) onto one surface of thesponge, and allowing 10 minutes for the solution to soak into thesponge. The sponges were held at a 90° angle, and lightly shaken toremove water not absorbed into the sponge. Sponges were weighed todetermine the volume of water which did absorb.

Solid discs were formed from various polymers in the PLGA family usingcompression molding. These discs were used to measure the contact angleof the polymers with water, and to test the ability of hepatocytes toadhere to the polymer films. Solid polymer discs were utilized in thesestudies instead of porous three-dimensional sponges as it simplified theanalysis (e.g., it is simpler to measure the advancing water contactangle on a two-dimensional film than in a three dimensional sponge).

Discs (0.5 mm thick) were formed from 0.75 g of PLA, polyglycolic acid,poly-D,L-lactic acid, or a 85/15 or 50/50 copolymer of lactic andglycolic acid (all purchased from Medisorb) using a Carver LaboratoryPress (Fred S. Carver, Inc.; Menominee Falls, Wis.). The polymer washeated to 185° C., and compressed at 1500 psi. Discs were coated withPVA or Pluronic F108 as described above. Contact angle measurements weremade from an advancing water droplet using a goniometer (Rame-Hart,Inc.; Mountain Lakes, N.J.). Reported values represent the mean andstandard error of the mean (SEM) calculated from the mean advancingcontact angle of a minimum of three films at each condition. The meanadvancing contact angle for each film was calculated from a minimum ofthree measurements.

B Results

Highly porous sponges were formed from PLA and 85/15 PLGA using apreviously described technique (Mikos et al., 1993b). The size and shapeof devices formed in this manner can be controlled by the geometry ofthe Teflon mold and the mass of salt particles and polymer loaded intothe mold. The porosity and pore size of devices formed with this type ofparticulate leaching technique can be controlled by varying the ratio ofpolymer/salt particles and the size of the salt particles (Mikos et al.,1993b). In this study, the ratio of polymer/salt was varied from 0.06 to0.03 to increase the device porosity from 90.±.1 to 95.±.1.5%, andcylindrical devices 1-5 mm thick (d=2 cm) were fabricated by increasingthe mass of polymer and salt from 0.412 to 2.06 grams.

Polymers of the lactic/glycolic acid family are all relativelyhydrophobic, as indicated by high contact angles with water. The contactangle of films fabricated from poly-L, lactic acid, polyglycolic acid,poly-D,L lactic acid, and 85/15 and 50/50 copolymers of lactic andglycolic acid were 79.degree±2°, 73.degree±2°, 72.degree±1.degree,73.degree±2°, and 69.degree±3°, respectively. To determine whether thehydrophobicity of these polymers could be decreased, solid films of PLAwere coated with the hydrophilic polymer PVA or a surfactant, Pluronic F108.

The advancing water contact angle decreased from 79.degree±2° to23.degree±2° when devices were exposed to PBS containing 10 mg/mLsolution of PVA. When PVA solutions ranging from 1 to 100 mg/mL wereused for coating, a similar decrease in the contact angle of polymerfilms was noted. The contact angle of PLA films also decreased from79.degree±2° to 22.degree±.3° when a 1 mg/mL solution of the Pluronicsurfactant was utilized in place of the PVA.

Contact angle decreases in the same range were found when devicesfabricated from PGA, or copolymers of lactic and glycolic acid weresimilarly treated with PVA (e.g., the contact angle of films of a 50/50copolymer decreased to 21.degree±7°). The decrease in the hydrophobicityof the polymer films was not permanent, however, as immersion of coatedpolymer discs into a PBS solution led to a rebound in the contact angleover a 24 hr period. For example, the contact angle of Pluronic-coateddevices returned to 76.degree±2° after 24 hr in PBS. These resultssuggest that the coating molecule re-dissolved over time in an aqueousenvironment.

Porous sponges fabricated from PLA were subsequently treated withaqueous solutions of PVA to determine whether PVA infiltration wouldimprove the ability of aqueous solutions to adsorb intothree-dimensional sponges. The sponge weight increased from 8 to 98% asthe PVA concentration was raised from 1 to 100 mg/mL. Infiltratingsponges with solutions of 1-10 mg/mL had a minimal effect on the devicepore size and porosity, but infiltrating with a 100 mg/mL solutionsignificantly decreased both. Devices infiltrated with a solutioncontaining 10 mg/mL of PVA rapidly and reproducibly absorbed aqueoussolutions equivalent to 98.±.1% of their pore volume, while untreateddevices only absorbed a volume of water equivalent to 6.±.2% of theirpore volume. Importantly, the porosity and pore size of devices returnedto their original values following exposure to an aqueous environmentfor 24 hr. Re-dissolution of the PVA infiltrating the sponge occurredduring this time, as the device weight returned to within 5.±.3% of theoriginal value under these conditions (similar results were obtainedwhen sponges fabricated from 85/15 PLGA were tested).

This method for fabricating devices has wide applicability. A variety ofother types of water soluble molecules, such as the Pluronicsurfactants, can also be utilized as the coating molecule. This type oftreatment will be useful for improving cell seeding into a variety ofhydrophobic polymer devices, both biodegradable and non-biodegradable.

EXAMPLE VI Mineral Patterning of Tissue Scaffolds with CaPO₄ with andwithout Fluoride

A Homogeneous Surface Mineralization

Pre-treatment to produce homogeneous surface hydrolysis may be achievedby either soaking in a NaOH solution or by treating with electromagnetic(EM) radiation as previously described. The treated biomaterial isincubated in a mineral-rich, preferably a calcium-rich, fluid, such as abody fluid or synthetic media that mimics body fluid, to spur nucleationand growth of a homogeneous mineral film on the surface. The fluidcontains about 10 mM CaF₂ where Fluoride is to be added to the tissuescaffold.

Functionalization and concomitant mineralization can also be achieved bysimply soaking in mineral-containing aqueous solutions, preferably inbody fluids or synthetic media that mimic body fluids. Preparation ofthe polymer biomaterials using a gas-foaming/particulate leachingprocess is generally preferred for such one step mineralization.

B. NaOH Pre-Treatment for Surface Mineralized Films

PLGA films (about 25 μm thickness) were prepared by a pressure castingtechnique. Raw polymer pellets were loaded into a mold and placed in aconvection oven at 200 degrees C. The molds were heated under pressure(about 22 N) for 30 sec and then cooled to room temperature.

For the creation of surface functional groups by NaOH treatment, thefilms were cleansed and immersed in 1.0 N NaOH solution for varyingtimes, up to 10 minutes to create surface functional groups. Followingimmersion, samples were rinsed 3.times in distilled water.

C. UV Pre-Treatment for Surface Mineralized Films

PLGA films (about 25 μm thickness) were prepared by a pressure castingtechnique. Raw polymer pellets were loaded into a mold and placed in aconvection oven at 200 degrees C. The molds were heated under pressure(about.22 N) for 30 sec and then cooled to room temperature.

For the creation of surface functional groups by UV (ultra violet)treatment, membranes were exposed to up to 8 hrs of surface irradiation.

D. Surface Mineralization after Pre-Treatment

Membranes treated by either NaOH treatment or UV treatment weresubsequently incubated at 37 degrees C. in 50 ml of a simulatedphysiological fluid (SPF, Na: 142 mM, K: 5 mM, Ca: 2.5 mM, Mg: 1.5 mM,Cl: 148 mM, HCO3: 4.2 mM, HPO₄: 1 mM, S04: 0.5 mm) buffered to pH 7.4.When fluoride was included, the solution contains 1.25 mM CaF₂.Solutions were replaced every 48 hours to ensure that there weresufficient ions in solution to induce mineral nucleation and growth.Following immersion for periods of 120 to 240 hours, samples were dried.

Fourier transform infrared (FTIR) analysis indicates the presence of asurface amorphous apatite. FTIR spectra of scaffolds treated for 0, 2,6, 10, and 16 days indicate the growth of a carbonated apatite mineralwithin the scaffold. Equivalent spectra were also produced with theUV-treated films. The broad band at 3570 cm⁻¹ is indicative of thestretching vibration of hydroxyl ions in absorbed water. The peak at1454 cm⁻¹ is indicative of CO₃ ²,.η3, while the 867 cm⁻¹ represents CO₃²,.η2. The peaks at 1097 cm⁻¹ and 555 cm⁻¹ are indicative ofanti-symmetric stretch (η3) and anti-symmetric bending (η4) of P0₄ ³,respectively. The peak at 1382 cm⁻¹ represents a NO₃ band.

The presence of OH⁻, C0₃ ²⁻ and PO₄ ³⁻ all indicate that an apatiticlayer has been formed. Other bands representative of apatites are maskedbecause of the strong absorption of the PLGA.

The major peaks at 1755 cm⁻¹ and 1423 cm⁻¹ represent PLGA, and the peakat 1134 cm⁻¹ indicative of C—O stretch in the ester. The peaks at 756cm⁻¹ and 956 cm⁻¹ are indicative of the amorphous domains of thepolymer.

The scaffolds demonstrated an increase in mass over time, culminating ina 11.±.2% mass gain at the end of the 16 day incubation. ANOVA ofpercent mass changes of experimental scaffolds reveal a significantdifference in scaffold mass over time (p<0.05), while ANOVA of percentmass changes of control scaffolds does not show a significant differenceover time (p>0.05). Percent mass changes of experimental samples andcontrol samples were significantly different for each time point beyond8 days (p<0.05).

To confirm that the increase in mass was caused by deposition of anapatitic mineral, the mass of phosphate in the scaffolds was nextanalyzed. Phosphate content within the treated scaffolds also increasedsignificantly with the treatment time. Comparison of phosphate massesvia ANOVA show a statistically significant increase over time (p<0.05),and the differences in phosphate mass between day 8 and 12 (p<0.05) andbetween day 12 and 14 (p=0.05) were also statistically significant.After a 14 day incubation, estimation of the mass of mineral on thescaffold using phosphate mass data gives 0.76 mg of hydroxyapatite,while the measured mass increase of the scaffold is 1.02.±.0.40 mg. Thefact that the measured value is larger than the estimated value suggestssignificant carbonate substitution in the mineral crystal.

Growth of the BLM layer significantly increased the compressive modulusof 85:15 PLG scaffolds without a significant decrease in scaffoldporosity. The compressive modulus increased from 60.±.20 KPa beforetreatment to 320.±.60 KPa after a 16 day treatment, a 5-fold increase inmodulus. ANOVA of modulus changes of experimental scaffolds reveal asignificant difference in scaffold modulus over time (p<0.05), whileANOVA of control modulus data does not show a significant differenceover time (p>0.05). The differences between moduli of experimentalscaffolds and moduli of control scaffolds were statistically significantfor treatment times of 10 days or longer (p<0.05). The porosity of thescaffolds did not decrease appreciably after incubation in SBF.Untreated scaffolds were 95.6.±.O.2% porous, while scaffolds incubatedin SBF for 16 days were 94.0.±.0.30% porous (n=3). This agrees with theelectron micrographs, which displayed only a thin (1-10 μm) mineralcoating, and thus no significant change in pore size due to mineralgrowth.

E One Step Mineralization

One step, room temperature incubation processes can also be used tocause nucleation and growth of mineral layers on polymer surfaces. Thisis achieved by incubating polymer scaffolds in mineral-containingaqueous solutions, such as body fluids and synthetic media that mimicbody fluids. These processes are able to grow bone-like minerals withinpolymer scaffolds in surprisingly simple and inexpensive methods. Theeffectiveness of these methods under room temperature conditions rendersthem conducive to the inclusion of bioactive proteins and othermaterials into the processing mineralization.

A first example of one step mineralization concerns the mineraldeposition on porous poly(lactide-co-glycolide) sponges via incubationin a simulated body fluid. The simple incubation technique was used toobtain nucleation and growth of a continuous carbonated apatite mineralon the interior pore surfaces of a porous, degradable polymer scaffold.Fluoride was also incorporated into the scaffold. The incorporation ofcalcium phosphate and fluoride into the scaffolds renders them suitablehydrogels to facilitate regeneration of enamel. In cases where no tissuein growth is needed, a hydrogel having a porosity of less than 50% orless than 40% or less than 30%, or less than 20%, or less than 10%, orbetween about 2% and about 10% may be used as a hydrogel plug in contactwith the enamel of a tooth.

A 3-dimensional, porous scaffold of 85:15 PLG was fabricated by asolvent casting/particulate leaching process and incubated in simulatedbody fluid (SBF; NaCl-141 mM, KCl-4.0 mM, MgSO₄-0.5 mM, MgCl₂-1.0 mM,NaHCO₃-4.2 mM, CaCl₂-2.5 mM, and KH2 PO₄-1.0 mM, with or without 1.25 mMCaF₂ in deionized H₂O, buffered to pH=7.4 with Trisma-HCl). Fouriertransform IR spectroscopy and SEM analyses after different incubationtimes demonstrated the growth of a continuous bone-like apatite layerwithin pores of the polymer scaffold.

The majority of the mineral growth occurred between days 8 and 12.Mineral growth into a continuous layer likely occurs from day 12, and iscomplete at or before day 16. The mineral grown, being continuous, isthus similar to that in bones and teeth.

The scaffolds demonstrated an increase in mass over time, with an11.±.2% gain after 16 days. The increase in mass is due to deposition ofan apatitic material. Quantification of phosphate on the scaffoldrevealed the growth and development of the mineral film over time withan incorporation of 0.43 mg of phosphate (equivalent to 0.76 mg ofhydroxyapatite) per scaffold after 14 days in SBF. The measured overallmass increase of the scaffold was 1.02.±.0.4 mg at 14 days. Thissuggests carbonate substitution in the mineral crystal.

The compressive moduli of polymer scaffolds also increased fivefold withformation of a mineral film after a 16 day incubation time, as opposedto control scaffolds. This was achieved without a significant decreasein scaffold porosity. The thin mineral coating is thus functionallyimportant, yet mineralization does not change the pore size.

85:15 PLGA scaffolds prepared by gas foaming/particulate leachingexhibit even more rapid nucleation and growth of apatitic mineral. The85:15 PLG scaffolds prepared via solvent casting/particulate leachingshowed a 3.±.1% increase in mass after a 6 day incubation in SBF. Incomparison, 85:15 PLGA scaffolds prepared by gas foaming/particulateleaching showed a mass increase of 6.±.1% after a 4 day incubation inSBF.

The even more rapid nucleation and growth of apatitic mineral on 85:15PLG scaffolds prepared by gas foaming/particulate leaching is believedto be due to the increase in carboxylic acid groups caused by the gasfoaming/particulate leaching process, i.e., the greater surfacefunctionalization. Leaching with 0.1 M CaCl₂ also likely facilitateschelation of Ca²⁺ ions, producing more rapid bone-like mineralnucleation.

F. Diffraction Lithography

Previous studies on the control of locations of cell adhesion to abiomaterial surface have utilized conventional UV lithography to patterna two dimensional polymer surface (Pierschbacher & Ruoslahti, 1984);Ruoslahti & Pierschbacher, 1987); Matsuda et al., 1990); Britland etal., 1992); Dulcey et al., 1991); Lom et al., 1993); Lopez et al.,1993); Healy et al., 1996).

In the prior techniques, the two-dimensional biomaterial surface iscoated with a thin layer of photoresist (PR), the PR is exposed througha metal mask, and the exposed PR is removed in solvent, leaving a PRmask on the surface of the biomaterial sample. The surface of thepolymer biomaterial is then chemically or physically treated through thePR mask, and the mask is removed by a solvent after treatment.

The former processes requires a flat, two dimensional biomaterial, whichsuffices for studying the effects of surface treatment on cell activity,but is not sufficient for the treatment of typical biomaterials, whichhave three dimensional surface contours.

In the present methods, suitable for use with three dimensionalpolymers, the grating produces a pattern of constructive and destructiveinterference on the polymer surface. As the grating is not required tobe in near contact with the biomaterial during treatment, thisdiffraction lithography process can be used to treat materials withcomplex three-dimensional surface contours. However, the process isequally useful in connection with two dimensional biomaterials.

G. Infiltration with Bioactive Molecules

Polymer biomaterial is treated to form a patterned biosurface,preferably using either patterned EM radiation or electron beamirradiation. Treated biomaterial is washed with distilled water toremove residual monomers from the surface photolysis or electrolysis.

The treated biomaterial is incubated in a solution containing bioactivemolecules or proteins, such as growth factors, adhesion molecules,cytokines and such like, which promote adhesion of a specific cell type.Cells are seeded onto the biomaterial in vitro in a cell culture medium.In vivo, cells attach to the biomaterial when implanted. In either case,cells adhere preferentially to the treated portions of the substrate.

The use of specific agents or proteins, such as growth factors, thatpromote attachment of certain cell types, gives the potential to patternany cell type on the three dimensional surface of the polymer, both invitro and in vivo.

EXAMPLE VII Making Tissue Scaffold Sponge Wafers Using Salt Frames

In this example, NaCl frames and biodegradable polymer scaffolds areproduced. The salt particles (Mallinkrodt, Paris, Ky.) were sieved toyield a range of sizes. NaCl crystals of a diameter of about 250-425 μmwere used.

Porous scaffolds were prepared either by solvent casting/particulateleaching, or gas foaming/particulate leaching processes using NaCl asthe particulate porogen. The solvent cast scaffolds were preparedessentially as described by Mikes, A. G., et al. (“Preparation andcharacterization of poly(L-Iactic acid) foams” Polymer 35:1068, 1994;which is incorporated herein by reference). NaCl molds were made bysubjecting NaCl crystals (diameter of about 250□425 μm) to 95% humidityfor periods from 0□24 hr to achieve fusion of NaCl crystals prior tosolvent casting. A closed, water-jacketed cell culture incubator (Form aScientific, Inc.) held at 37° C. was used to create a 95% humidityenvironment for fusion of NaCl crystals.

Poly(lactide-co-glycolide) (PLG) pellets with a lactide:glycolide ratioof 85:15 were obtained from Medisorb, Inc. (intrinsic viscosity(I.V.)=0.78 dl/g) and Boehringer-Ingelheim Inc. (I.V.=1.5 dl/g). Highinherent viscosity PLG was used in the solvent casting process to ensurethat the scaffolds would retain adequate mechanical integrity despitetheir relatively high porosity (˜97%). PLG pellets were dissolved inchloroform (Mallinkrodt, Paris, Ky.) to yield a solution of 10%weight/volume (w/v). The polymer solution was then poured into an NaClcontaining mold wherein the salt crystals had been fused, as describedabove. Following solvent evaporation, the salt was removed by immersionin distilled water for about 48 hours.

The gas foamed scaffolds were essentially prepared as described byHarris, L. D., et al. (“Open pore biodegradable matrices formed with gasfoaming” J Biomed Mater Res 42:396, 1998; which is incorporated hereinby reference). NaCl molds were made by subjecting NaCl crystals(diameter of about 250□425 μm) to 95% humidity for periods from 0□24 hrto achieve fusion of NaCl crystals prior to solvent casting. Followingtreatment in 95% humidity samples were dried in a vacuum desiccator for48 hr before further processing. A closed, water-jacketed cell cultureincubator (Form a Scientific, Inc.) held at 370 C was used to create a95% humidity environment for fusion of NaCl crystals. PLG pellets(prepared as above) were dissolved in chloroform. Frames of fused NaClwere mixed with PLG were loaded into an aluminum die (1.35 cm diameter;Aldrich Chemical Co., Milwaukee, Wis.) and was compressed at 1500 Psifor 1 minute using a Carver Laboratory Press (Fred S. Carver, Inc.,Menominee Falls, Wis.) to yield solid disks (thickness of about 3.4 mm).The samples were then exposed to high pressure CO₂ gas (800 psi) for 24hours to saturate the polymer with gas. A thermodynamic instability thenwas created by decreasing the gas pressure to ambient pressure. Thislead to the nucleation and growth of CO₂ pores within the polymermatrices. The NaCl particles subsequently were removed from the matricesby leaching the matrices in distilled water for 48 hours. All processingsteps were performed at ambient temperature.

Scaffolds were circular disks with a diameter of about 12 mm and athickness of about 3 mm. The pore size range was controlled by usingNaCl particles with a diameter of about 250□425 μm in the processing.The total porosity of scaffolds was calculated using the known densityof the solid polymer, the measured polymer mass of the scaffold, and themeasured external volume of the scaffold.

Incubation of NaCl crystals in 95% humidity resulted in fusion of thecrystals, creating a highly interconnected NaCl matrix. Fused salt moldswere bisected and imaged prior to solvent casting to observe the extentof NaCl crystal fusion. In addition, polymer scaffolds were bisectedafter preparation via freeze fracture. A carbon coating was evaporatedonto the surface of each bisected salt mold and polymer scaffold, andsamples were imaged under high vacuum using a Hitachi S□3200N SEMoperating at 20030 kV. Fusion of salt crystals prior to addition of PLGin chloroform (solvent casting) resulted in enhanced poreinterconnectivity within the scaffold. The pore structure within thescaffolds appears similar to the structure of the fused salt matrix, asexpected. Pores within the cross section of 1 hr salt fusion (SF)samples display a defined pore structure with intermittent holes in porewalls, while the cross section of scaffolds created from 24 hr SFsamples display a much less organized pore structure and a very largedensity of holes in pore walls. The hole size increased significantlywith fusion time, from an average diameter of 31+10 μm after 1 hour offusion to 78+21 μm after 24 hours of fusion (p<0.05). In addition, thepore walls in the 24 hr SF scaffolds display thickness contours suchthat the walls appear thicker in the area adjacent to the holes in porewalls and along the outer diameter of the walls. A higher magnificationview of a pore wall within a 24 hr SF scaffold further displays thecontoured structure of the pore walls. The salt fusion process had noeffect on the porosity of the scaffolds, and the calculated totalporosities of the solvent cast scaffolds for each salt fusion timeperiod were 97+1%.

A close examination of the electron micrographs of the solvent castscaffolds formed after 1 and 24 hours of NaCl fusion indicate that theexposure to 95% humidity has caused several important changes in thestructure of the salt particles. In addition to the formation of bridgesbetween particles at the points of contact, the radius of curvature ofedges and corners in individual particles of salt has increased. Thesechanges are shown schematically. The radius of curvature of saltcrystals was calculated from electron micrographs. The pixel size foreach image was calibrated, and the pencil tool was used to mark tangentpoints on crystal edges. The calibration values and pixel coordinateswere then used to calculate the cord length between tangent points,which was multiplied by (2/2 to obtain the crystal radius of curvature.The diameter of holes in pore walls was determined by measuring themajor and minor diametral axes of each hole using microsoft paint andtaking the average. The increased radius of curvature at the edges andcorners of each particle of salt results in an increased sphericity ofeach particle, and thus in each resulting pore in the scaffold. The meanradius of curvature of the crystal edges increased from 19+10 μm, to32+15 μm after 12 hours of exposure to 95% humidity, then to 62+18, μmafter a full 24 hours of exposure. As a result, many of the smallercrystals became nearly spherical in shape after 24 hours of fusion. Oneadditional consequence is that thicker polymeric struts may be formed inthe space vacated by the corners and edges of each salt crystal, whichmay result in the thickness contours in pore walls described above andin varied mechanical properties.

Fusion of salt crystals in PLG/NaCl pellets prior to gas foaming alsoresulted in a pronounced variation in pore structure. The cross sectionof 1 hr SF samples shows small holes in pore walls similar to those inthe solvent cast 1 hr SF samples. The 24 hr salt fusion samples lack adefined pore structure and pores appear to simply feed into each other.The gas foamed SF scaffolds do not display any of the contours in porewalls observed in the solvent cast SF samples. Again, the salt fusionprocess had no effect on the total scaffold porosity. The totalporosities of the gas foamed scaffolds for each salt fusion time periodwere 94+1%.

Fusion of salt crystals for 24 hr resulted in a 2 fold increase in thecompressive modulus of the solvent cast scaffolds. Compressive moduli ofscaffolds were determined using an MTS Bionix 100 mechanical testingsystem. Samples were compressed between platens with a constantdeformation rate of 1 mm/min. Compression plates had a diameter of 45mm, and thus covered the entire 12 mm diameter surface of the scaffold.A small pre-load was applied to each sample to ensure that the entirescaffold surface was in contact with the compression plates prior totesting, and the distance between plates prior to each test was equal tothe measured thickness of the scaffold being tested. Compressive moduliwere determined for scaffolds without salt fusion and for each of foursamples for each salt fusion time. Values on graphs represent means andstandard deviations. Statistical analysis was performed using InStatsoftware, version 2.01. At each time point, experimental moduli werecompared to control moduli via a Student's t-test to reveal significantdifferences in compressive modulus. No significant modulus change isobserved after 1 hr, or 12 hr of salt fusion. Alternatively, there was astatistically significant decrease in the compressive modulus of gasfoamed scaffolds processed using salt fusion when compared with controlscaffolds.

EXAMPLE VIII Incorporation of Morphogenic Agents into a Tissue ScaffoldWafer (HYDROGEL?)

A. Wafer Preparation and Characterization

Wafers containing BMP-7 are prepared by a modification of a previouslydescribed double-emulsion technique (Cohen et al., 1991). In brief, a75/125 copolymer of poly-(D,L-lactic-co-glycolic) acid (Resomer RG 75R,intrinsic viscosity 0.2; Henley Chem. Inc., Montvale, N.J.) is dissolvedin ethyl acetate (Fisher Scientific) to yield a 5% solution (w:v).Recombinant human BMP 7 is dissolved in water to yield a solution of 2mg/ml, and 50 ml of the BMP-7 solution is added to 1 ml of the polymersolution. The polymer/BMP-7 solution is sonicated continuously at 10watts (Vibracell; Sonics and Materials, Danbury, Conn.) for 15 sec toyield a single emulsion.

An equal volume of an aqueous solution containing 1% polyvinyl alcohol(MW 25,000, 88% hydrolyzed; Polysciences Inc., Warrington, Pa.) and 7%ethyl acetate is added to the single emulsion, and the resultingsolution is vortexed (Vortex Mixer; VWR) for 15 sec at the high settingto yield the double emulsion. This double emulsion is transferred to arapidly stirring 250 ml beaker containing 150 ml of an aqueous solutionof 0.3% polyvinyl alcohol/7% ethyl acetate. The double emulsion isdistributed into casting molds of selected dimensions for the wafer, andthe ethyl acetate is allowed to evaporate over the ensuing 3 hr to yieldpolymer wafers with entrapped BMP-7. The wafers are then filtered andwashed with water. The wafers are lyophilized (Labconco Freeze Dryer,Kansas City, Mo.), and stored at −20° C. until use. Control wafers areprepared with the same procedure, but the aqueous solution used to formthe first, single emulsion (water in organic) contained no BMP-7.

To determine the efficiency of BMP-7 incorporation and the kinetics ofBMP-7 release from the wafers, the BMP-7 is labelled with ¹²⁵I usingstandard labelling techniques to obtain a specific activity of about 10to 1000 mCi/mg. Approximately 1 μCi of labelled BMP-7 is added to theaqueous BMP-7 solution before formation of the single emulsion, and thewafers are prepared as described above. After wafer fabrication, a knownmass of wafers is counted in a scintillation counter and theincorporated cpm is compared to that of the initial aqueous BMP-7solution to calculate the percentage of the total BMP-7 that isincorporated into the wafers.

To determine the release of BMP-7 from wafers, a known mass of wafers(approximately 10 mg) prepared with the labelled EGF is placed in aknown volume (2 ml) of phosphate buffered saline (PBS) solutioncontaining 0.1% Tween 20, (Sigma Chem. Co.) and placed in an incubatormaintained at 37° C. At set times, the solution is centrifuged toconcentrate the beads at the bottom of the vial, and samples (0.1 ml) ofthe PBS/Tween 20 solution is removed. The sample volume is replaced withfresh PBS/Tween 20 solution. The amount of ¹²⁵I-BMP-7 released from thewafers is determined (n=4) at each time point by counting the removedsample in a gamma counter, and compared to the ¹²⁵I-BMP-7 loaded intothe wafers. The maximum theoretical BMP-7 concentration in the releasemedium (approximately 5 mg/ml) should be well below the maximumsolubility of BMP-7, thus establishing sink conditions for the releasestudy.

B. Results

The yield of wafers with this process should be about 92+5%.

To determine the efficiency of BMP-7 incorporation into wafers, and therelease profile from the same the ¹²⁵I-labelled BMP-7 is utilized as atracer. Approximately 1/2 of the initial BMP-7 (53+11%) will beincorporated into wafers. When BMP-7-containing wafers are placed in anaqueous medium, an initial burst of BMP-7 release will be noted. Afterthis time BMP-7 will be released in a steady manner over the remainderof a 30 day time course.

The time over which a drug is released from a polymer matrix cantypically be regulated by the drug loading, the type of polymerutilized, and the exact processing conditions (Mooney et al., 1992). Therelease of protein from copolymers of lactic and glycolic acid, such asutilized in this Example, is generally controlled by the erosion of thepolymer when the protein/polymer ratio is low (Cohen et al., 1991). Thereleased protein must retain its biological activity for this approachto be useful. The biological activity of the BMP-7 incorporated into andreleased from wafers in this Example will not be adversely affected.This approach to delivering BMP-7 can also be readily expanded todeliver other molecules such as BMP-2, BMP-4. VEGF or any morphogenicprotein, alone or in combination with such proteins.

EXAMPLE IX Alginate Hydrogel Growth Factor Incorporation and Releasefrom a Foamed Matrix Tissue Scaffold

About 1 to 100 mg of ¹²⁵I-labelled protein growth factor is first addedto a solution of 1% sodium alginate, and then beads of this solution aregelled by injecting droplets into an aqueous solution containing calciumchloride. The alginate beads (approximately 3 mm in diameter) arecollected, rinsed, and lyophilized. The lyophilized beads are mixed with85:15 PLGA and NaCl particles and the mixture is compression moulded andprocessed with the gas foaming/particulate leaching process aspreviously described.

To test for release, after salt leaching and drying, the matrices areplaced in serum free tissue culture medium and maintained at 37° C.Medium samples are taken periodically, and analyzed for the content of1251-VEGF (released from PLGA matrices). The released growth factor maybe normalized to the total incorporated growth factor.

Results will show that an initial burst of approximately 20% of theincorporated growth factor occurs in the first day, and a sustainedrelease of growth factor will be noted for at 20 days.

EXAMPLE X Growth Factor Incorporation and Release from MineralizedMatrices

A. Materials and Methods

1. Gas Foaming-Particulate Leaching

Poly(lactide-co-glycolide) pellets with a lactide:glycolide ratio of85:15 obtained from Medisorb, Inc. (I.V.=0.78 dl/g) and ground to aparticle size between 106 and 250 μm. Ground PLGA particles are thencombined with 250 μl of a 1% alginate (MVM, ProNova; Oslo, Norway)solution in ddH.₂O, with about 1 to about 1000 μg of a morphogenicprotein such as BMP-2, BMP-4, BMP-7 or VEGF and the like, orcombinations of such morphogenic proteins. These solutions arelyophilized, mixed with 100 mg of NaCl particles (250 μm<d<425 μm), andcompression moulded at 1500 psi for 1 min in a die of suitable size. Atypical 4.2 mm diameter die yields 2.8 mm thick disks with a diameter of4.2 mm.

Disks are then exposed to 850 psi CO₂ gas in an isolated pressure vesseland allowed to equilibrate for 20 h. The pressure is decreased toambient in 2 min, causing thermodynamic instability, and subsequentformation of gas pores in the polymer particles. The polymer particlesexpand and conglomerate to form a continuous scaffold with entrappedalginate, morphogenic protein, and NaCl particles. After gas foaming,the disks are incubated in 0.1 M CaCl₂ for 24 h to leach out the saltparticles and induce gelation of the alginate within the polymer matrix.Alginate is included in the scaffolds because it has been shown to abatethe release of VEGF from PLGA scaffolds (Wheeler et al., 1998).

2. Mineralization

Certain scaffolds are mineralized via a 5 day incubation in a simulatedbody fluid (SBF). Simulated body fluid (SBF) was prepared by dissolvingthe following reagents in deionized H₂O: NaCl-141 mM, KCl-4.0 mM,MgSO₄-0.5 mM, MgCl₂-1.0 mM, NaHCO₃-4.2 mM, CaCl₂-2.5 mM, and KH2 PO₄-1.0mM. The resulting SBF is buffered to pH 7.4 with Trisma-HCl and held at37.degree C. during the incubation periods. The SBF solutions arerefreshed daily to ensure adequate ionic concentrations for mineralgrowth.

The porosity of scaffolds is calculated before and after mineralizationtreatment using the known density of the solid polymer, the knowndensity of carbonated apatite, the measured mass of mineral and polymerin the scaffolds, and the volume of the scaffold.

3. Characterization of Mineral Growth

To analyze mineral growth on gas foamed PLG scaffolds, sets of threescaffolds were incubated in SBF for periods ranging from 0-10 days.Samples were removed from solution and analyzed after 0, 2, 4, 8, and 10day incubation periods. The dry mass of each scaffold is measured beforeand after incubation in SBF, and percent increases in mass is calculatedand compared using ANOVA and a Student's t-test to reveal significantdifferences in mass for different SBF incubation times.

The amount of phosphate present in the scaffolds after theaforementioned incubation times is determined using a previouslydescribed colorimetric assay (Murphy et al., J. Biomed. Mat. Res., InPress; incorporated herein by reference). The phosphate mass data werealso compared using ANOVA and a Student's t-test to reveal significantdifferences in mass for different SBF incubation times.

To estimate the amount of apatite on the scaffold after a 6 dayincubation, the measured mass of phosphate is multiplied by the knownratio of mass of hydroxyapatite [Ca₁₀(PO₄)₆ (OH)₂, f.w.=1004.36 g] tomass of phosphate in hydroxyapatite (569.58 g). This is a conservativeestimate, since it assumes that all phosphate is being incorporated intostoichiometric hydroxyapatite. This mineral mass estimate increases ifone assumes increasing substitution of carbonate into the mineralcrystal.

4. Growth Factor Release Measurements

In order to assess the incorporation efficiency of the morphogenicgrowth factors into the PLG scaffolds and to track the growth factorrelease kinetics from the scaffolds, the growth factor is labelled with¹²⁵I to a specific activity of about 1 to 1000.μCi/.μg in place of theunlabelled growth factor in the normal sample preparation. To assessgrowth factor incorporation efficiency, the total incorporated activityis compared to the activity of the initial ¹²⁵I growth factor sampleprior to incorporation into the scaffolds.

To determine the effects of mineral growth on factor release, releasekinetics are measured both in SBF during mineral formation and inphosphate buffered saline (PBS). Scaffolds prepared with radiolabeledgrowth factor are placed in 4 ml of SBF or PBS and held at 37 degrees C.At various set times, the scaffolds are removed from solution and theirradioactivity is assessed using a gamma counter. After each analysis,solutions are refreshed and scaffolds are placed back into solution.

The amount of radiolabeled growth factor released from the scaffolds isdetermined at each time point by comparing the remaining ¹²⁵I growthfactor to the total originally loaded into each scaffold. The percentrelease of VEGF from scaffolds incubated in SBF is compared to that ofscaffolds incubated in PBS at each time point via a Student's t-test toreveal significant differences in cumulative release.

B. Results

1. Mineralization

Incubation of gas foamed 85:15 poly(lactide-co-glycolide) scaffoldscontaining growth factors results in the growth of bone-like mineral onthe inner pore surfaces. Analysis of variance will show that differencesin percent mass gain with SBF incubation time is significant (p<0.05).The scaffolds show an increase in mass with incubation time, with a 6±1%mass gain after a 4 day incubation in SBF. The scaffold mass willsubsequently remain relatively constant. The increase in mass betweentwo day and four day incubation times will be significant p<0.05), whilethere will be no significant difference in percent mass gain between thefour day incubation time and the longer incubation times (p>0.05).

To verify that the increase in mass is caused by the deposition of anapatitic mineral, the mass of phosphate in the scaffolds can beanalyzed. Phosphate content within scaffolds increased with SBFincubation time. Analysis of variance will show that differences inphosphate content with SBF incubation time is significant (p<0.05). Thedifference in phosphate content between the two day and six dayincubation times will be significant (p<0.05), while there will be nosignificant difference between the phosphate mass of the six dayincubation time and longer incubation times (p>0.05).

It has previously been shown that the increase in mass and phosphatecontent in these scaffolds indicates growth of a continuous bone-likemineral film on the inner pore surfaces (Murphy et al., J. Biomed. Mat.Res., In Press).

The total porosity of the scaffolds after a 10 day incubation in SBF isabout 92±1%, which is similar to the initial scaffold porosity (93±1%).

After a 6 day incubation, estimation of the mass of mineral on thescaffold using phosphate mass data gives 0.10 mg of hydroxyapatite,while the measured mass increase of the scaffold is 0.39.±.0.03 mg. Thefact that the measured value is larger than the estimated valueindicates significant carbonate substitution in the mineral crystal.

2. Growth Factor Release and Activity

Growth factors are incorporated into PLGA scaffolds with an efficiencyof 44±9% and is released over a 15 day period in SBF and PBS solutions.An initial burst release of the incorporated growth factor will beobserved over the first 12-36 h followed by a sustained release for 20days or more.

The cumulative release from scaffolds incubated in SBF becomessignificantly smaller than release from scaffolds incubated in PBS after3 days, and this difference will remain significant through 10 days ofrelease (p<0.05). At time points beyond 10 days there is no significantdifference in cumulative release from scaffolds incubated in SBF versusthose incubated in PBS.

The foregoing examples are included to demonstrate preparation orexecution of preferred embodiments of the invention. It should beappreciated by those of skill in the art that the techniques disclosedin the examples that follow represent techniques discovered by theinventors to function well in the practice of the invention, and thuscan be considered to constitute preferred modes for its practice.However, those of skill in the art should, in light of the presentdisclosure, appreciate that many changes can be made in the specificembodiments that are disclosed and still obtain a like or similar resultwithout departing from the spirit and scope of the invention.

1. A composition for treating dental tissue, comprising a tissuescaffold wafer comprising a scaffolding material associated with calciumphosphate and fluoride.
 2. The composition of claim 1 further comprisinga physiologically effective amount of a morphogenic agent that promotesgrowth of dentin tissue.
 3. The composition of claim 2 wherein themorphogenic agent is encoded by a member of the TGF-β supergene family.4. The composition of claim 2 wherein the morphogenic agent is selectedfrom the group consisting of BMP-2, BMP 4, BMP-7, VEGF, FGF-1, FGF-2,IGF-1, IGF-2, PDGF, GDF-1, GDF-2, GDF-3, GDF-4, and GDF-5.
 5. Thecomposition of claim 2 wherein the morphogenic agent is selected fromthe group consisting of BMP-2, BMP 4, BMP-7, and GDF-5.
 6. Thecomposition of claim 1 further comprising an active agent selected fromthe group consisting of an anti-bacterial agent and an anti-inflammatoryagent.
 7. The composition of claim 1 wherein the tissue scaffold iscomprised of scaffolding polymer selected from the group consisting ofPLLA, PDLLA, PGA and PLGA.
 8. The composition of claim 1 wherein thetissue scaffold is comprised of PLGA.
 9. A vacuum manipulator formanipulating a tissue scaffold wafer, comprising: a vacuum tube having aproximal end, a distal end, and walls between the proximal and distalends enclosing the vacuum tube; an attachment at the distal end of thevacuum tube to permit fluid communication between a vacuum source andthe vacuum tube; a suction cup attached to the proximal end of thevacuum tube in fluid communication with the vacuum tube, the suction cupbeing sized to fit onto a surface of a wafer comprised of dentalscaffolding material; and a valve assembly positioned at the distal endof the vacuum tube proximate to the vacuum source, the valve assemblybeing operable to close and open fluid access between the vacuum tubeand a vacuum source and to open and close fluid access between thevacuum tube and a pressure source.
 10. The vacuum manipulator of claim 9wherein the valve assembly is manipulated by a manual dial.
 11. Thevacuum manipulator of claim 9 further including an in-line filterdisposed in the vacuum tube between the suction cup and the distal end,the in-line filter having sufficient pore volume to allow passage of thefluid between the suction cup and the vacuum tube while preventing thepassage of a pathogen.
 12. The vacuum manipulator of claim 9 wherein thevacuum tube has a bend along the length thereof to orient the proximalend for insertion into a mouth of human along a plane defined by crownsof teeth.
 13. The vacuum manipulator of claim 9 wherein the vacuumsource is comprised of a pliable bulb shaped member.
 14. The vacuummanipulator of claim 9 wherein the attachment for the vacuum source iscomprised of a hose attachable to an external vacuum source.
 15. Thevacuum manipulator of claim 9 further comprising a source of positivepressure in fluid communication with the vacuum tube through the vacuumassembly.
 16. The vacuum manipulator of claim 15 wherein the positivepressure source is ambient air pressure.
 17. The vacuum manipulator ofclaim 15 wherein the pressure source is a source of positive air flowthat is greater than ambient air pressure.
 18. The vacuum manipulator ofclaim 9 wherein the valve assembly is adjustable to allow a controlledamount of at least one of positive or negative pressure to be drawn orapplied within the vacuum tube.